Silicone Surface with Drug Nanodepots for Medical Devices - ACS

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Silicone surface with drug nano-depots for medical devices Jiratchaya Mokkaphan, Wijit Banlunara, Tanapat Palaga, Premsuda Sombuntham, and Supason Pattanaargson Wanichwecharungruang ACS Appl. Mater. Interfaces, Just Accepted Manuscript • DOI: 10.1021/am505566m • Publication Date (Web): 14 Oct 2014 Downloaded from http://pubs.acs.org on October 19, 2014

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ACS Applied Materials & Interfaces

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Silicone surface with drug nano-depots for medical

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devices

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Jiratchaya Mokkaphan, a Wijit Banlunara, b Tanapat Palaga, c,f Premsuda Sombuntham d and

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Supason Wanichwecharungruang e,f,* a

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Program of Petrochemical and Polymer Science, Faculty of Science, b Department of

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Pathology, Faculty of Veterinary Science, c Department of Microbiology, Faculty of Science, d

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Department of Otolaryngology Head and Neck surgery, Faculty of Medicine, e Department of

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Chemistry, Faculty of Science, Chulalongkorn University,

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f

Nanotec-CU Center of Excellence on Food and Agriculture, Bangkok 10330, Thailand

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TITLE RUNNING HEAD: Silicone surface with drug nano-depots

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KEYWORDS: silicone surface, drug release, surface functionalization, inflammation,

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implantation, endotracheal tube

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ACS Paragon Plus Environment

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ABSTRACT: Ideal surface of poly(dimethyl siloxane) (PDMS) medical devices requires

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sustained drug release to combat with various tissue responses and infection. At present, non-

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covalent surface coating with drug molecules using binders possesses detachment problem, while

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covalently linking drug molecules to the surface provides no releasable drug. Here, a platform

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that allows the deposition of diverse drugs onto the PDMS surface in an adequate quantity with

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reliable attachment and a sustained release character is demonstrated. First, a PDMS surface with

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carboxyl functionality (PDMS-COOH) is generated by subjecting a PDMS piece to an oxygen

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plasma treatment to obtain silanol moieties on its surface, then condensing the silanols with 3-

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aminopropyltriethoxysilane molecules to generate amino groups, and finally reacting the amino

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groups with succinic anhydride. The drug-loaded carriers with hydroxyl groups on their surface

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can then be esterified to the PDMS-COOH, resulting in PDMS surface covalently grafted with

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drug-filled nanocarriers so that the drugs inside the securely grafted carriers can be released.

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Demonstrated here is the covalently linking the surface of PDMS endotracheal tube with

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budesonide-loaded ethyl cellulose (EC) nanoparticles. A secure and high drug accumulation at

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the surface of the tubes (0.025 mg/cm2) can be achieved without changing its bulk property such

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as hardness (Shore-A), and sustained release of budesonide with a high release flux during the

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first week followed by a reduced release flux over the subsequent three weeks can be obtained.

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In addition, the grafted tube possesses more hydrophilic surface, thus is more tissue-compatible.

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The grafted PDMS pieces show a reduced in vitro inflammation in cell culture, and a lower level

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of in vivo tissue responses, including a reduced level of inflammation, compared to the

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unmodified PDMS pieces, when implanted in rats. Although demonstrated with budesonide and

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a PDMS endotracheal tube, this platform of grafting a PDMS surface with drug-loaded particles

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can be applied to other drugs and other devices.


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INTRODUCTION

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Due to its ease of fabrication (ease of cross-linking), biocompatibility and advantageous

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chemical (bio-inertness, oxidative and thermal stability) and physical (optical transparency and

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adjustable modulus) characteristics, poly(dimethyl siloxane) or PDMS is a popular thermosetting

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silicone elastomer used for the production of medical devices to be put inside the body, either

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temporarily or permanently, such as catheters, artificial prosthetics and subcutaneous sensors.

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Nevertheless, inflammation and infection occur frequently at the tissues surrounding the

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devices.1–4 The hydrophobic nature of the PDMS surface contributes greatly to the formation of a

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dead-space between the hydrophilic tissue and the implanted devices, resulting in the

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colonization of microorganisms and the formation of biofilms.5,6 Moreover, the low surface free

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energy of PDMS imparts a poor wettability to the material that results in tissue irritation,

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abrasion and ulceration.7,8 In fact, the surface wettability directly influences the interaction

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between tissue cells and surfaces and is, therefore, one of the first concerns when developing

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new devices for biomedical applications.9 This factor is critically important for percutaneous

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devices which pass through the skin from the outside to the interior of the body as bacteria can

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enter the body along the devices surface.

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The problems of the hydrophobic surface of PDMS together with its frequent infection

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and inflammation after being implanted are required to be solved without altering the excellent

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bulk properties of the material. Various surface functionalization strategies have been carried out

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to make the PDMS surface more hydrophilic or even charged, but infection and inflammation

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still occur.10-13 Coating the surface of PDMS medical devices with drug molecules, using various

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non-covalent interactions and binders, to give the devices a drug release character,14 has resulted


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in peel-off and drug migration problems with no significant improvement compared to the

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uncoated devices.15 Covalent linking of drug molecules to the surface is another approach to

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lessen the infection, inflammation and tissue reaction levels.16-18 However, in addition to a low

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drug content on the surface, another limitation of this approach is that the tethered drug moieties

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are fixed at the device’s surface and can neither migrate into nor build up to a high drug

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concentration in the surrounding tissue despite the fact that a high drug concentration during the

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critical post-implantation period is required for a better improvement of the post-implantation

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reaction.16,19 Moreover, this approach requires different drugs to be individually customized onto

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the surface of the devices. Installing a drug release character to any PDMS materials has also

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been carried out by directly dispersing drug molecules into a polymeric matrix.20,21 Nevertheless,

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this direct mixing method is only popular in pharmaceutical preparations and devices, such as a

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silicone elastomer vaginal gel with antiviral activity,22 intra-vaginal drug delivery rings23-25 and

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transdermal drug delivery patches,26 in which mechanical properties and long-term stability of

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the devices are not critical. To conserve mechanical and viscoelastic properties of PDMS

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medical devices, only limited amounts of drugs could be mixed into the polymeric matrix, thus a

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release of sufficient amounts of drugs could not be attained through this strategy.14

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Therefore, a platform that allows the deposition of diverse drugs onto the device’s surface in

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an adequate quantity with reliable attachment and a sustained release character is needed. Here, a

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novel platform for covalently linking nano-depots filled with a high amount of drug molecules to

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the surface of the PDMS was developed for the first time. The platform offers the secure

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attachment and ability to release drug molecules into the surrounding tissue, as well as an

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increased hydrophilicity of the PDMS surface, without altering the key physical characters of the

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bulk PDMS material.


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Endotracheal tubes for assisted ventilation are a high-risk medical device in terms of

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infection and inflammation due to their location and function.1 As a result, the tracheal tube was

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chosen as a representative PDMS medical device with budesonide, a hydrophobic drug with an

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anti-inflammatory activity, as a representative drug to be deposited on to the tube surface. Here

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the surface functionalization of the PDMS tracheostomy tube, to make it both hydrophilic and

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able to slowly release budesonide, was developed by loading budesonide into ethyl cellulose

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(EC) nanoparticles and then covalently linking the budesonide-EC nanoparticles onto the

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functionalized PDMS tube. The process was confirmed by scanning electron microcopy (SEM)

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and confocal fluorescent microscopy (CLFM). The budesonide loading level was evaluated

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along with its subsequent in vitro release over a 4-week period. The physical property of the

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(budesonide-loaded particle)-grafted PDMS tubes was compared to that of the untreated PDMS

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tubes in terms of their hardness (Shore-A) and surface hydrophilic nature. The obtained

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(budesonide-loaded particle)-grafted PDMS tubes were tested for their in vitro and in vivo anti-

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inflammatory activity in cell culture and after implanting into rats, respectively.

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EXPERIMENTAL SECTION

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The PDMS tubes with a 10.0 mm outer diameter and 8.00 mm inner diameter were

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obtained from Koken (Tokyo, Japan), whilst EC (48% (w/w) ethoxy content, viscosity of 300 cp)

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and 1-hydroxybenzotriazole (HOBt) were from Sigma Aldrich (Steinheim, Germany).

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Budesonide (MW 430.5, analytical grade) was purchased from Fluka (Seelze, Germany).

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Succinic anhydride (analytical grade), 3-aminopropyltriethoxysilane (γ-APS) and 1-ethyl-3-(3-

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dimethylaminopropyl)carbodiimide-HCl (EDCI) were purchased from Acros Organics (Geel,

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Belgium). Rhodamine B sulfonyl chloride was purchased from Invitrogen (Grand Island, NY,

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USA).


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UV–Vis spectroscopy was obtained from a UV-2550 spectrometer (SHIMADZU, Japan).

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The SEM and transmission electron microscopy (TEM) were performed using a JEM-6400

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(JEOL, Japan) and JEM-2100 (JEOL, Japan) instrument, respectively. Dynamic light scattering

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(DLS) was performed on a Zetasizer nano series instrument (Zs, Malvern Instruments, USA).

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The water contact angle was obtained using a Ramé-Hart Model 200, Standard Contact Angle

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Goniometer (Succasunna, NJ, USA). Fluorescent images were obtained by CLFM using a Nikon

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Digital Eclipse C1si Confocal Microscope system (Tokyo, Japan) equipped with Plan

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Apochromat VC 100×, Melles Griot Diode Laser and 85 YCA-series Laser at 561 nm, a Nikon

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TE2000-U microscope, a 32-channel-PMT-spectral-detector and Nikon-EZ-C1 Gold Version

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3.80 software. Hardness (Shore-A) of the materials was measured using a Durometer Model 408

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Type A (Shore A) (PTC Instruments, Los Angeles, CA, USA). 1H Nuclear magnetic resonance

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(NMR) spectra were obtained in deuterated chloroform (CDCl3-d1) with tetramethylsilane

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(TMS) as an internal reference using ACF 200 spectrometer which operated at 400 MHz for 1H

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nuclei (Varian Company, Palo Alto, CA, USA).

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Encapsulation of budesonide into EC nanoparticles. The encapsulation of budesonide into EC

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nanoparticles was performed at a 1:1 (w/w) ratio of budesonide: EC by solvent displacement.27

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Budesonide (500 mg) and EC (500 mg) were dissolved in ethanol to give a 5 mg/mL final

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concentration of each of EC and budesonide. The mixture (50 mL) was dialyzed against distilled

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water (5 × 1000 mL, using CelluSep T4 membrane, MW cut-off (MWCO) of 12–14 kDa,

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Membrane Filtration Products, Seguin, TX, USA). An aqueous suspension of budesonide-loaded

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EC particles obtained inside the dialysis bag was collected. The morphology and size of the dry

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budesonide-loaded EC particles were evaluated by SEM and TEM, whilst the hydrated size was

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determined by DLS. The amounts of budesonide in the particles and those in the dialysate water


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were determined using UV/VIS absorption spectroscopy at 246 nm, with the aid of a calibration

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curve constructed with 5, 10, 15, 20, 25, 30 and 35 ppm standard budesonide solutions. The

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encapsulation efficiency (EE) and loading capacity were calculated from equations (1) and (2),

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respectively:

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EE (%) = (A/B) × 100,

(1)

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Loading capacity (%) = (A/C) × 100,

(2)

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where A is the weight of encapsulated budesonide, B is the weight of budesonide originally used

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and C is the weight of the obtained budesonide-loaded particles.

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The obtained budesonide-loaded EC nanoparticle suspension was then freeze-dried.

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Fluorescence labeled particles. The fluorophore Rhodamine B was covalently linked to the EC

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particles by adding 0.5 mL of a 1% (w/v) Rhodamine B into 200 mL of an aqueous EC

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nanoparticle suspension (2.5 mg/mL) in the presence of a catalytic amount (2 drops) of pyridine

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under a light-proof and nitrogen (N2) atmosphere (Scheme S1 in supporting information or SI).

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The mixture was stirred at room temperature overnight. The Rhodamine B-labeled EC polymer

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was then purified from free Rhodamine B by dialysis under light-proof condition.

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Surface treatment of the PDMS tubes (Scheme 1). The PDMS tubes were first subjected to

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oxygen-plasma treatment (STERRAD100S, ASP, Irvine, CA, USA) at 55–60 °C with radio

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wave in 18 mL of 59% (v/v) hydrogen peroxide (H2O2) for 10 cycles. The cycle was composed

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of sequential vacuum (397 mtorr, 20 min 16 sec), injection (6.34 torr, 6 min 2 sec), diffusion (15

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torr, 2 min), plasma (552 mtorr, 7 min 23 sec), injection (7.67 torr, 6 min 2 sec), diffusion (15

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torr, 2min), plasma (630 mtorr, 7 min 11 sec) and vent stages. The hydrophilic/hydrophobic

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nature of the obtained tubes was then evaluated by measuring the water contact angle.


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Condensation between the silanol groups on the surface of the plasma-treated PDMS

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tubes with γ-APS was then performed using a method modified from that previously described.28

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Briefly, the plasma-treated PDMS tubes (130 cm2 total surface area) were cut into 1 cm pieces

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and then refluxed in 50 mL of 2% (v/v) γ-APS in toluene with 2-3 drops of triethanolamine

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(TEA), for 3 h under a N2 atmosphere. The tubes were then removed from the solution and

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sequentially rinsed with water (1000 mL), hexane (40 mL) and ethanol (40 mL), and annealed at

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room temperature for 24 h. The washed tubes were dried in a desiccator for 24 h to obtain the

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amino-functionalized PDMS (PDMS-NH2) tubes. The product was characterized for the

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presence of surface primary amino groups by the Kaiser test.29

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The PDMS-NH2 tubes (130 cm2 total surface area) were then stirred in a solution of

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succinic anhydride (1 g) in 50 mL dimethylformamide (DMF) with 2-3 drops of pyridine at 70

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°C overnight under a N2 atmosphere. Free succinic anhydride and succinic acid were eliminated

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by washing the obtained tubes with excess water. The obtained carboxyl PDMS (PDMS-COOH)

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tubes were then subjected to the Kaiser test.

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For the Kaiser test, the tubes were soaked in 0.1% (w/v) ninhydrin solution in EtOH and

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then heated at 50–60 °C for 1 min to activate the reaction. The color change was visually

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observed.

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Grafting the surface of PDMS-COOH tube with budesonide-loaded EC particles (Scheme 1).

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To have the surface of PDMS-COOH grafted with the budesonide-loaded EC particles, a 50 mL

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aqueous suspension of budesonide-loaded EC nanoparticles (containing 125 mg budesonide and

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250 mg EC) was put into a round-bottomed flask and HOBt (0.88 g, 0.65 mol) in ethanol (5 mL)

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was added. The PDMS-COOH tubes (4.40 cm2 surface area per tube, 30 tubes) and EDCI (1.25

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g, 0.65 mol) were added into the flask and the mixture was stirred at 0 ºC under a N2 atmosphere


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overnight. The grafted tubes (denoted as budesonide-loaded PDMS tubes) were then rinsed with

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excess water repeatedly followed by 20 mL of ethanol and then dried before being subjected to

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SEM and hardness (Shore A) analyses.

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PDMS-COOH tubes were also grafted with fluorescent EC particles using a similar procedure

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and the obtained fluorescent particle-grafted tube was subjected to CLFM analysis (excitation at

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561 nm and detection by scanning from 565–700 nm).

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The amount of budesonide deposited onto the tubes was quantified by subjecting the tubes to

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ethanol extraction. The ethanolic extract was centrifugally filtered through a membrane (MWCO

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of 10 kDa) to filter out the EC and the supernatant was subjected to HPLC analysis with the aid

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of a calibration curve. The extracted budesonide was also dried, re-dissolved in CDCl3 and

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subjected to 1H NMR analysis.

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Scheme 1. Grafting of budesonide-loaded EC particles onto the surface of PDMS tube.


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In vitro budesonide release. The budesonide-loaded PDMS tubes (seven tubes, 4.40 cm2

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surface area per tube) were submerged in 5.0 mL of release medium (0.075 mM phosphate

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buffer saline solution (pH 5.8), 5% (v/v) Tween 20 and 20% EtOH) at 37 °C. Aliquots of the

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release medium (500 µL each) were transferred into micro centrifuge tubes at the indicated times

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(0, 1, 2, 5, 8, 9, 10, 12 and 14 d). The volume of the release medium was kept constant by adding

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fresh release medium (500 µL each) after each sampling. The withdrawn aliquot was gently

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preconcentrated under a N2 gas flow to a final volume of ~200–300 µL, whereupon 200 µL of

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ethyl acetate was added, the mixture was vortexed, allowed to phase separate and the ethyl

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acetate layer was collected. The budesonide concentration in the ethyl acetate layer was then

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quantiﬁed using high performance liquid chromatography (HPLC) analysis with the aid of a

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calibration curve constructed from budesonide standard solutions freshly prepared in ethyl

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acetate.

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HPLC analysis. The HPLC analysis was performed using a Waters 1525 binary HPLC (pump)

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and a 100 mm × 4.6 mm column packed with Hypersil C18 (Thermo Fisher Inc, Waltham, MA,

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USA) and connected to a Waters 2489 UV-Vis detector (Milford, MA, USA) operating at 240

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nm. The sample injection volume was 10 µL and the mobile phase was a 2:30:68 (v/v/v) ethanol:

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acetonitrile: 25 mM phosphate buffer (pH 3) mixture, and the flow rate was 1.5 mL/min.

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In vitro anti-inflammatory activity in cell culture. The unmodified PDMS (control) and the

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budesonide-loaded PDMS tubes were subjected to an in vitro anti-inflammatory test in tissue

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culture using the RAW264.7 (ATCC TIB-71) mouse macrophage-like cell line. The tubes were

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cut into 0.5 × 0.6 cm pieces (2 mm thick). For the budesonide-loaded PDMS pieces, both sides

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were grafted with budesonide-loaded EC particles and contained 15 µg of budenoside/piece. The


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tubes were autoclaved prior to use. These were evaluated along with a culture control (no test

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sample added) and free budesonide.

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Cells were maintained in complete medium (CM; Dulbecco's Modified Eagle Medium

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supplemented with 10% (v/v) fetal bovine serum, 1% (w/v) sodium pyruvate, 1% (w/v) HEPES

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(Hyclone), 100 U/ml penicillin and 0.25 mg/ml streptomycin) and were incubated at 37 ˚C in a

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humidified 5% (v/v) CO2 incubator. Cells were plated overnight in a 24-well plate (5 x 104

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cells/well) in CM. The medium was changed to fresh CM before the addition of the respective

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treatment (nothing, budenoside (15 µg), unmodified PDMS or budesonide-loaded PDMS pieces)

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for 48 h. The media was then replaced with fresh CM and the RAW264.7 cells were stimulated

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with 100 ng/ml LPS (from Salmonella) and 10 ng/mL interferon-γ (IFNγ) for 24 h. The culture

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supernatant was then harvested and the concentration of nitric oxide (NO) was measured using

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the Griess reaction as described previously.30

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In vivo anti-inflammatory activity in rat implants. The procedures were approved by both the

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ethics committee and the institutional animal care and use committee (IACUC) of Chulalongkorn

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University. The unmodified PDMS and budesonide-loaded PDMS pieces were subjected to an in

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vivo anti-inflammatory test by subcutaneous implantation in rats, according to the biological

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evaluation of medical devices of international standard ISO10993-6:2007 (E). The PDMS and

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budesonide-loaded PDMS pieces were prepared as described above, except that each piece was

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1.0 x 0.6 cm in size and so contained 30 µg of budenoside/piece (for the budesonide-loaded

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PDMS pieces). The autoclaved pieces were used in the experiment.

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In total nine male Wistar rats, aged 12 weeks old (National Laboratory Animal Center, NLAC,

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Nakorn Pathom, Thailand) were surgically implanted with the PDMS pieces on one side of the

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dorsal subcutaneous tissue of the back and the budesonide-loaded PDMS pieces on the other side


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(sides chosen at random) under ketamine hydrochloride and xylazine hydrochloride general

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anesthesia. Each three randomized animals were sacrificed and the skin samples removed after 7,

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14 and 28 days post-surgery, respectively. The skin samples were fixed in 10% (w/v) buffered

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formalin and histopathologically examined by hematoxylin and eosin (H&E) staining. Semi-

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quantitative evaluation of the degree of inflammation was performed by histopathological

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grading according to the ISO10993-6:2007(E) Annex E.

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RESULTS AND DISCUSSION

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Budesonide-loaded carriers. Budesonide was successfully encapsulated into the EC

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nanoparticles by solvent displacement. The aqueous suspension of the budesonide-loaded EC

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nanoparticles formed inside the dialysis bag appeared as a milky white suspension (Figure 1).

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The SEM and TEM images of the product (Figure 1) revealed a spherical morphology with an

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anhydrous particle size of 342.10 ± 85.04 nm while DLS gave a slightly larger hydrodynamic

248

diameter of 356 nm with a polydispersity index of 0.256 (mid-range polydispersity). The mean

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zeta potential of the budesonide-loaded EC nanoparticles in water at pH 6.5 was -57.40 ± 1.27

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mV, agreeing well with the observed stable particle suspension over 3 months.

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Figure 1. Budesonide-loaded EC particles: (A) The aqueous suspension of budesonide-loaded

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EC nanoparticles (concentrations of budesonide and EC of 1.25 and 2.50 mg/mL, respectively),

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and representative (B) SEM and (C) TEM images of the budesonide-loaded EC nanoparticles.


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Centrifugal filtration of the particle suspension allowed the separation of the

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unencapsulated budesonide (filtrate) from the budesonide-loaded EC particles (residue), and the

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filtrate was quantified for budesonide concentration by UV absorption spectrophotometry with

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the aid of a calibration curve (see Figure S1 in SI for calibration curve). The result indicated that

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the obtained budesonide-loaded EC nanoparticles had a high loading capacity (33.26 ± 0.23%

260

(w/w)) and the encapsulation process gave a moderate EE (49.83 ± 0.11%). It should be noted

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here that the encapsulation process was performed at a 1:1 (w/w) ratio of budesonide: EC. The

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optimized condition used here gave particles with no budesonide crystal/precipitate.

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During dialysis of the ethanolic solution containing EC and budesonide, when ethanol was

264

slowly displaced with water, it was likely that the EC chains self-assembled themselves in such a

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way that the hydrophilic hydroxyl groups were in maximum contact with water medium while

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the hydrophobic ethoxyl groups were at the inside of the forming particles away from water

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medium, resulting in budesonide-loaded EC particulates with a hydrophilic outer surface and a

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hydrophobic interior.27 During the particle formation, the hydrophobic nature of budesonide

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probably drove the molecules to the hydrophobic core of the forming spheres, leading to

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encapsulation of budesonide inside the forming EC particles.

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The fact that the particles were loaded with the hydrophobic budesonide molecules at 33%

272

loading content, agreed well with this molecular assembling speculation. In addition, an

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observation that a fast addition of water into an EC solution (in ethanol) produced water un-

274

dispersible precipitates while the slow displacement of ethanol in the EC solution with water

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through the dialysis resulted in water dispersible nanospheres, also conformed well with the

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proposed self-assembling speculation.


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Surface functionalization of PDMS tubes (Scheme 1). To make PDMS tubes that were able to

278

slowly release budesonide while still maintaining their bulk properties, only the surface of the

279

tubes was covalently linked to the budesonide-loaded EC nanoparticles. The process was

280

comprised of the four steps of: (i) activation of the PDMS surface with oxygen plasma to

281

introduce silanol moieties onto the surface (plasma-treated PDMS), (ii) silanization of the

282

plasma-treated PDMS surface with γ-APS to introduce amino groups onto the tubes’ surface

283

(PDMS-NH2), (iii) coupling the amino groups with succinic anhydride to introduce carboxylic

284

groups onto the tubes’ surface (PDMS-COOH), and (iv) linking the hydroxyl groups on the outer

285

surface of the budesonide-loaded EC particles to the carboxylic groups on the surface of the

286

PDMS-COOH tubes through ester bonds (budesonide-loaded PDMS), as shown in Scheme 1.

287

Since the surface of PDMS tubes is quite inert, activation of the surface by oxygen

288

plasma was first performed to introduce silanol (Si-OH) moieties onto the surface via a free

289

radical oxidation mechanism. This treatment rendered the PDMS surface hydrophilic, which was

290

verified by the change in the water contact angle from 107.28 ± 1.9 ° for the untreated PDMS

291

surface to 71.6 ± 6.5 ° for the plasma treated surface. Note that the water contact angle was

292

measured soon after the plasma treatment was completed because, theoretically, the resulting

293

silanol groups on the plasma-treated PDMS could undergo crosslinking into stable siloxane

294

linkages. The plasma-treated PDMS tubes were then quickly silanized by reaction with γ-APS to

295

introduce amino groups onto their surface via a base catalyzed hydrolysis-condensation between

296

the silanol groups and the γ-APS to yield the covalent siloxane network with the aminopropyl

297

moieties on the outer surface of the tubes (PDMS-NH2).31

298

Attempts to verify the chemical change at the surface of the plasma-treated tubes with

299

attenuated total reflectance Fourier transform infrared spectroscopy (ATR-FTIR) analysis failed


14

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300

because the IR spectrum of the freshly prepared plasma-treated PDMS and that of the untreated

301

PDMS tubes were the same. This was likely because the functionalization was only attained at

302

the top surface of the material, but the ATR-FTIR detection acquired information of the sample

303

at a depth of more than 60–70 nm from the surface. Attempts to scratch the functionalized

304

surface off and powder it so as to then make a KBr compression disk also failed since the PDMS

305

was too elastomeric to scratch. Nevertheless, in addition to the water contact angle information,

306

the successful coupling with γ-APS confirmed the presence of the silanol groups on the surface

307

of the plasma-treated PDMS tubes.

308

A presence of the amino functionality on the surface of the thoroughly washed and

309

annealed tubes was confirmed by coloration with the Kaiser reagent,29 in which a deep violet

310

blue color on the tube surface was clearly observed (Figure S2 in SI). The obtained amino-

311

functionalized tubes were stable and could be kept in a desiccator until needed.

312

Here, the plan was to covalently link the budesonide-loaded EC particles to the surface of

313

the PDMS tubes so the budesonide-loaded EC particles could act as drug depots at the tubes’

314

surface. Many of the hydroxyl groups on the EC chains should be exposed at the outer surface of

315

the budesonide-loaded EC particles and so could be used for covalently linking to the

316

functionalized PDMS. Thus, carboxylic groups were introduced onto the surface of the tubes to

317

serve as a reactive functionality for linking with the hydroxyl groups on the EC particles via ester

318

linkage. Accordingly the PDMS-NH2 tubes were reacted with succinic anhydride to produce the

319

PDMS-COOH tubes (Scheme 1). Successful coupling was verified through the negative Kaiser

320

test (Figure S3 in SI).

321

Finally, the PDMS-COOH tubes were grafted with the budesonide-loaded EC particles

322

through the formation of ester linkages between the surface hydroxyl and carboxyl groups of the


15


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Page 16 of 39

323

particles and the tubes, respectively (Scheme 1). The SEM images of the obtained budesonide-

324

loaded PDMS tubes clearly indicated the presence of particles on the surface of the tube (Figure

325

2 A), and their size resembled that of the budesonide-loaded EC particles (Figure 1). To more

326

clearly characterize the distribution of the particles on the tube’s surface, the EC particles were

327

first covalently labeled with Rhodamine B, a fluorescent chromophore, and then the PDMS-

328

COOH tubes were grafted with the fluorescent-EC particles using a similar process as above.

329

The obtained (fluorescent EC particle)-grafted PDMS tubes showed an obvious fluorescence

330

signal on their surfaces with an island-like distribution of particles on the tube’s surface (Figure

331

2 B). Factors affecting the distribution of particles on the surface probably included the

332

distribution of –COOH functionality on the surface (started from the plasma treatment step) and

333

also agitation during the reaction between PDMS-COOH and the particles. Nevertheless,

334

improved hydrophilic nature of the grafted tubes was observed, i.e., the budesonide-loaded tubes

335

showed a better wetting with a water contact angle of 48.6 ± 8.5 ° compared to 107.28 ± 1.9 ° for

336

the unmodified PDMS tubes. Overall, the data support that the tube’s surface was grafted with

337

the particles. The size and thickness of the tubes did not change upon functionalization, but the

338

surface of the budesonide-loaded PDMS tubes appeared more opaque compared to the

339

unmodified tubes (Figure 2 C-D).

340

Quantification of the budesonide level on budesonide-loaded PDMS tubes using ethanol

341

extraction followed by quantification with HPLC, revealed that, under the grafting condition

342

used, the average budesonide coverage on the tube’s surface was 0.0254 ± 0.022 mg/cm2. Here

343

HPLC was used to quantify budesonide in the extract, instead of the previously used direct UV

344

absorption analysis, in order to make sure that the UV absorption detected at 240 nm was only

345

from budesonide, not from other components that might be extracted out from the tubes. Note


16

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346

here that under the HPLC condition used, retention time of budesonide was approximately 11

347

min and linear calibration curve was obtained for the 10-50 ppm budesonide standards (Figure

348

S4, SI). Since the extracted budesonide retained its chemical structure (verified by 1H NMR), it

349

was concluded that no harm was done to the drug molecules during the grafting process and so

350

the PDMS surface could be non-covalently loaded with budesonide by covalently linking the

351

budesonide-loaded EC particles to its surface. This allows the accumulation of a high drug

352

content on the limited surface area. More importantly, the accumulated drug molecules were not

353

destroyed by the reaction that linked the particles to the PDMS surface. These grafted particles

354

should act as drug depots on the PDMS surface to release active drug molecules in situ during

355

use, and this was evaluated next.


17


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Page 18 of 39

356 357

Figure 2. Images of PDMS tubes. SEM (A) and fluorescence (B) image of the surface of the

358

(budesonide-loaded EC particle)-grafted PDMS tube. Photographs of (budesonide-loaded EC

359

particle)-grafted PDMS tubes (C) and original PDMS tubes (D).

360

To make sure that the grafting of the tube’s surface with budesonide-loaded EC

361

nanoparticles did not significantly alter any important physical properties of the tube, the

362

hardness (Shore-A) of the tube as a representative and key character was measured using a

363

durometer. Numerically similar values of 26.3 ± 2.1 and 28.1 ± 0.6 were obtained for the


18

Page 19 of 39

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364

unmodified and the budesonide-loaded PDMS tubes, respectively, and these were not

365

significantly different (P ≤ 0.05; Mann-Whitney U test). Note that a direct comparison of the

366

hardness between the budesonide-loaded PDMS and the unmodified PDMS tubes was possible

367

because they had the same thickness (2 mm). Thus, the surface treatment did not significantly

368

change the hardness, a key physical property, of the PDMS tubes.

369

Budesonide release. The in vitro release of budesonide from the budesonide-loaded PDMS

370

tubes was evaluated in the release buffer of pH 5.8. Initial fast release followed with slower

371

release during later time was observed (Figure 3), thus the release mechanism was a passive

372

diffusion driven by concentration gradient between the inside and the outside of the particles, in

373

which the EC matrix acted as a diffusion barrier. The fast release during the first week was a

374

result of the high drug concentration gradient between the inside and the outside of the grafted

375

particles. The fastest release during day one was expected since the concentration difference was

376

maximal. The fast release during the first few days fulfills the requirement of providing a high

377

drug concentration during the critical first few days post-implantation. Thereafter, the release

378

rate decreased as the concentration gradient decreased and became essentially constant after 1

379

week, which is probably controlled by the rate of diffusion through the EC matrix of the

380

budesonide molecules that was located deep inside the grafted particles. It should be noted here

381

that EC is stable at alkaline pH, down to pH value of approximately 4, therefore, the sustained

382

release character of the encapsulated drug from the tube’s surface could be expected in

383

applications at pH above 4. Nevertheless, an actual release character will be affected greatly by

384

the environment around the tube. Although the release of budesonide from the surface of the

385

budesonide-loaded PDMS tubes shown here did not exactly represent the true in situ (in vivo)

386

release at the contact tissue in a real application, the results still confirmed the sustained drug


19


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Page 20 of 39

387

release characteristic of the budesonide-loaded PDMS tubes. An ability to release budesonide

388

from the budesonide-loaded tubes was also tested in two more circumstances, in tissue culture

389

medium and in contact with rat’s tissue under real implantation, and the results and discussions

390

are presented next.

391 392

Figure 3. Release of budesonide from the budesonide-loaded PDMS tubes when immersed in

393

release buffer at 37 °C. Data are shown as mean ± 1 standard deviation and are derived from

394

three independent repeats.

395

In vitro anti-inflammatory activity in cell culture. Treatment of the macrophage-like cell line

396

RAW264.7 in vitro with the budesonide-loaded PDMS tubes in the presence of LPS and IFN-γ

397

resulted in a significantly lower nitric oxide production level than when treated with the

398

unmodified PDMS (Figure 4). This implied that the budesonide inside the grafted particles on

399

the tubes could be released into the culture medium and performed in vitro anti-inflammatory

400

action to the cells.

401


20

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402 403

Figure 4. In vitro anti-inflammatory activity (shown as a decreased NO production level) of the

404

budesonide-loaded PDMS when tested on the LPS and IFN-γ co-stimulated RAW264.7

405

macrophage cell line, in comparison with free budesonide (at the same amount of budesonide as

406

that in the budesonide-loaded PDMS sample) and the unmodified PDMS. Data are shown as the

407

mean ± 1 standard deviation and are derived from three independent repeats. Different letters on

408

the data bars indicate statistical significant difference between the samples.

409

3.5. In vivo anti-inflammatory activity in implanted rats

410

Evaluation of the tissue surrounding the subcutaneous implanted budesonide-loaded

411

PDMS and unmodified PDMS (control) tubes in nine male Wistar rats was performed visually

412

and by histopathological examination under a light microscope. Visual observation of the

413

implant area at day 7 post-implantation revealed no large differences in the tissue reactions in the

414

area surrounding either the unmodified PDMS or the budesonide-loaded PDMS tube implants,

415

the surgical reactions had masked any developing specific inflammation and tissue reactions. At

416

14, 21 and 28 days post-implantation, however, the area around the budesonide-loaded PDMS

417

implants showed significantly less inflammation.

418

Subcutaneous implantation of a medical device, usually creates a wound that triggers the

419

host inflammatory response and infiltration of phagocytic cells, such as polymorphonuclear,


21


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Page 22 of 39

420

lymphocytes, macrophages, giant and mast cells. Here, the amounts of these cells were observed

421

by histopathological examination via H&E staining. A significantly lower inflammatory cell

422

infiltration level and less tissue responses in the tissues surrounding the budesonide-loaded

423

PDMS implants were observed during the 28-day post-implantation period compared to those

424

surrounding the unmodified PDMS ones (Table 1). At the 7th and 14th day post-implantation all

425

the rats showed a significantly lower number of the five types of inflammatory cells in the area

426

surrounding the budesonide-loaded PDMS implant, compared to the control area in the same rat

427

in each case. On the 28th day post-implantation polymorphonuclear cells could not be observed

428

in the surrounding area of either the budesonide-loaded PDMS or control PDMS implants, but

429

fewer macrophages and lymphocytes were observed around the area implanted with the

430

budesonide-loaded PDMS tubes compared to that around the control PDMS implants. In

431

addition, no giant or mast cells were observed around the budesonide-loaded PDMS implants in

432

contrast to the control PDMS implants. The tissue responses examined, which included fibrin

433

clots, hemorrhage, fibrosis and neovascularization, were all significantly less pronounced in the

434

area surrounding the budesonide-loaded PDMS implants than the control PDMS implants in all

435

nine rats during the 28-day post-implantation period (Table 1). Representative H&E stained

436

tissue samples at the 28th day post-implantation are shown in Figure 5, with inflammatory cells

437

clearly identifiable by their dark purple colored nuclei. The tissue at the implant interface of the

438

control (unmodified PDMS) showed infiltration of a large number of inflammatory cells,

439

whereas that adjacent to the budesonide-loaded PDMS implants exhibited a markedly lower

440

inflammatory response as evidenced by fewer cells and a thinner fibrous tissue level.

441


22

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442

Table 1. Histopathological findings of the cellular and tissue responses in the tissues around the

443

implanted materials. Cell types *

Day

7

14

28

444

Tissue responses **

Polymorphonuclear cells

Lymphocytes

Macrophages

Giant cells

Mast Fibrin cells clot

Hemorrhage

Fibrosis

Neovascularization

Unmodified

2

3

3

1

1

2

1

1

1

Budesonide -loaded

1

1

2

0

0

1

0

1

1

Unmodified

1

2

3

1

2

0

0

2

3

Budesonide -loaded

0

1

1

0

0

0

0

1

1

Unmodified

0

2

3

1

1

0

0

3

2

Budesonide -loaded

0

1

1

0

0

0

0

2

1

PDMS types

Data shown are the average from the nine rats (three rats per group for each time point).

445 446 447

Scores: * Cell types: phf = per high power field (400 x); 0 = none, +1 = 1–5 cells/phf, +2 = 5– 10 cells/phf, +3 = heavy infiltrate, +4 = packed; graded from more than 9 slides obtained from each surgical implant wound (a total of more than 27 slides for each sample at each time point)

448 449

**Tissue responses: 0 = none, +1 = minimal, +2 = mild, +3 = moderate, +4 = severe; graded from 3 rats at each time point.

450 451

The lower inflammatory and tissue responses observed in the budesonide-loaded PDMS

452

implants compared to the unmodified PDMS ones were likely to have resulted from the

453

continuous release of budesonide, an anti-inflammatory drug, from the surface of the

454

budesonide-loaded PDMS implants into the surrounding tissues. In addition, the more

455

hydrophilic surface of the budesonide-loaded PDMS implant should reduce the irritation,

456

abrasion and dead space between the implant and the tissue7,8 and so contributed to the reduction


23


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Page 24 of 39

457

in the inflammation and tissue responses. Four key points should be noted for this particle-

458

grafted PDMS platform (illustrated here as budesonide-loaded PDMS). First, the covalent

459

linkage of the carrier particles (EC in this case) to the modified PDMS surface provides a secure

460

attachment of the drug carrier at the device’s surface, and so detachment during storage or use is

461

unlikely. This is different from previous reports that have relied on the non-covalent binding.14,32

462

Second, this platform gave devices with a high drug surface coverage without any deep alteration

463

to the bulk (PDMS) material. Thus, the hardness, a key physical property of the bulk platform,

464

remained unchanged. This is different from the direct mixing of drugs into polymer matrix in

465

which intense alterations of mechanical and viscoelastic properties and material homogeneity of

466

the devices could be observed even at low drug loading.14 Third, the drug molecules were not

467

covalently linked to the surface, but were intended to be released and diffuse away from the

468

immediate implant surface to be able to perform their bioactivity in the surrounding tissue (in

469

this case, the anti-inflammatory activity of budesonide as the payload drug). In other words, this

470

new platform provides a secure grafting, yet the drug molecules are still releasable. This is

471

different from the direct grafting of the drug molecules onto the surface where they remain

472

tethered without any active release16-18 and so exert no biological activity away from the implant

473

surface and would also be poor delivery agents for drugs requiring cellular internalization for full

474

bioactivity. The platform presented here provides a secure deposit of drugs for subsequent active

475

release from the PDMS surface. The flexibility of the platform, in terms of an easy change of the

476

payload drug (or even drugs), should allow a fast exploration for making other PDMS medical

477

devices with the active release of various drugs. Here, we used nanocarriers made from EC as a

478

model payload carrier, but carriers made of other materials should also be useable. Nevertheless,

479

nanocarriers made from EC are easy to fabricate and have a demonstrated ability to encapsulate


24

Page 25 of 39

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480

many hydrophobic drugs with a high loading capacity.27,33,34 Fourthly and lastly, the use of

481

nanoparticles with a hydrophilic outer surface also rendered the grafted PDMS surface

482

hydrophilic, and this is important for better tissue contact. This will also help reduce the problem

483

of the dead space between the implant and the tissue that is encountered with the unmodified

484

hydrophobic PDMS surface, and so minimize the bacterial colonization and abrasion. In fact, this

485

was observed in the implantations in rats in this study where lower tissue responses and less

486

inflammation were observed.

487 488

Figure 5. Histopathology of the H&E-stained subcutaneous tissue after 4 weeks post-

489

implantation. Optical micrographs (400 x magnification) showing the fibrous tissue surrounding

490

the (A) unmodified PDMS implant and (B) the budesonide-loaded PDMS implant. Nuclei of

491

inflammatory cells are stained dark purple while the collagen appears pink (H&E stain). The

492

implant was located at the bottom white region (denoted with arrow) of each image. Each image


25


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Page 26 of 39

493

shown here is a representative of more than nine images obtained from each of the three rats

494

(total of more than 27 slides).

495 496

CONCLUSION

497

Here silicone tracheostomy tubes with a sustained release of budesonide were prepared

498

by covalently grafting the surface of the PDMS tubes with the budesonide-loaded EC

499

nanoparticles. Firstly, silanol groups were introduced onto the surface of PDMS tubes by oxygen

500

plasma treatment, and then the plasma-treated PDMS tubes were subjected to condensation

501

reaction with γ-APS to yield PDMS tubes with amino groups on their surface (PDMS-NH2).

502

Coupling of the PDMS-NH2 with succinic anhydride then yielded tubes with carboxylic moieties

503

on the surface, and these PDMS-COOH tubes were then esterified with the hydroxyl groups on

504

the surface of the budesonide-loaded EC particles to yield the budesonide-loaded PDMS tubes. A

505

high budesonide-loading level of the attached spheres of 33.26% (w/w) was obtained, resulting

506

in a high budesonide coverage of 0.025 mg/cm2 at the tube’s surface. Importantly, since the

507

functionalization was restricted to only the surface of the tube, the hardness of the bulk PDMS

508

material was not significantly changed. Sustained in vitro release of the payload budesonide from

509

the budesonide-loaded PDMS tubes was observed for 4 weeks, and so the grafted budesonide-

510

loaded EC particles acted as nano-drug depots on the device surface. The budesonide-loaded

511

tubes could be kept at room temperature and autoclaved prior to use. Although the grafting of the

512

PDMS surface with the drug-loaded particles is shown here for a tracheostomy tube with

513

budesonide as the demonstrated drug and EC particle as the depot, the platform could

514

theoretically be applied to different PDMS medical devices and other payload drugs and to not

515

only indwelling devices but also to permanent implants requiring a controlled release of drug


26

Page 27 of 39

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516

during the critical initial post-implantation period, and so should be a significant improvement in

517

the success of implantations.

518 519

ASSOCIATED CONTENT

520

Labeling EC particles with Rhodmine B. Colorimetric Kaiser test for amine groups on the

521

surface of unmodified PDMS and PDMS-NH2 tubes. Colorimetric Kaiser test for amine groups

522

on the surface of the PDMS-NH2 tubes before and after treatment with succinimide. This

523

material is available free of charge via the Internet at http://pubs.acs.org.

524

AUTHOR INFORMATION

525

Corresponding Author

526

Supason

527

Wanichwecharungruang,

Department

of

Chemistry,

Faculty

of

Science,

Chulalongkorn University, Bangkok 10330, Thailand. E-mail: [email protected]

528

FUNDING SOURCES

529

This work was financially supported by the Ratchadapisek-sompot Endowment Fund of

530

Chulalongkorn University (RES560530097-Adv Mat) and Nanotec-CU Center of Excellence on

531

Food and Agriculture.

532

ABBREVIATIONS

533

PDMS, poly(dimethyl siloxane); PDMS-NH2, amino-fuctionalized poly(dimethyl siloxane);

534

PDMS-COOH, carboxyl functionalized poly(dimethyl siloxane); budesonide-loaded PDMS,

535

poly(dimethyl

536

aminopropyltriethoxysilane.

537

siloxane)

grafted

with

budesonide-loaded

particles;

γ-APS,

3-

REFERENCES


27


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538 539 540 541

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TOC

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Silicone surface with securely attached drug depots that can actively release drug molecules


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Scheme 1. Grafting of budesonide-loaded EC particles onto the surface of PDMS tube. Surface treatment of the PDMS tubes (Scheme 1). The PDMS tubes were first subjected to oxygen-plasma treatment (STERRAD100S, ASP, Irvine, CA, USA) at 55–60 °C with radio wave in 18 mL of 59% (v/v) hydrogen peroxide (H2O2) for 10 cycles. The cycle was composed of sequential vacuum (397 mtorr, 20 min 16 sec), injection (6.34 torr, 6 min 2 sec), diffusion (15 torr, 2 min), plasma (552 mtorr, 7 min 23 sec), injection (7.67 torr, 6 min 2 sec), diffusion (15 torr, 2min), plasma (630 mtorr, 7 min 11 sec) and vent stages. The hydrophilic/hydrophobic nature of the obtained tubes was then evaluated by measuring the water contact angle. 106x130mm (300 x 300 DPI)



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Figure 1. Budesonide-loaded EC particles: (A) The aqueous suspension of budesonide-loaded EC nanoparticles (concentrations of budesonide and EC of 1.25 and 2.50 mg/mL, respectively), and representative (B) SEM and (C) TEM images of the budesonide-loaded EC nanoparticles. 36x16mm (300 x 300 DPI)


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Figure 2. Images of PDMS tubes. SEM image of the surface of the (budesonide-loaded EC particle)-grafted PDMS tube (A). Fluorescence image of the surface of the (fluorescent EC particle)-grafted PDMS tube. Photographs of (budesonide-loaded EC particle)-grafted PDMS tubes (C) and original PDMS tubes (D). 149x272mm (300 x 300 DPI)



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Figure 3. Release of budesonide from the budesonide-loaded PDMS tubes when immersed in release buffer at 37 °C. Data are shown as mean ± 1 standard deviation and are derived from three independent repeats. 74x65mm (300 x 300 DPI)


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Figure 4. In vitro anti-inflammatory activity (shown as a decreased NO production level) of the budesonideloaded PDMS when tested on the LPS and IFN-γ co-stimulated RAW264.7 macrophage cell line, in comparison with free budesonide (at the same amount of budesonide as that in the budesonide-loaded PDMS sample) and the unmodified PDMS. Data are shown as the mean ± 1 standard deviation and are derived from three independent repeats. Different symbols above the data bars indicate significant statistical difference (P < 0.05) in comparison to other data bars. 51x32mm (300 x 300 DPI)



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Figure 5. Histopathology of the H&E-stained subcutaneous tissue after 4 weeks post-implantation. Optical micrographs (400 x magnification) showing the fibrous tissue surrounding the (A) unmodified PDMS implant and (B) the budesonide-loaded PDMS implant. Nuclei of inflammatory cells are stained dark purple while the collagen appears pink (H&E stain). The implant was located at the bottom white region (denoted with arrow) of each image. 109x144mm (300 x 300 DPI)


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Graphic for Table of Content 35x15mm (300 x 300 DPI)


Silicone Surface with Drug Nanodepots for Medical Devices - ACS

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