Stimulus-Responsive Antibiotic Releasing Systems for the Treatment

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Stimulus-Responsive Antibiotic Releasing Systems for the Treatment of Wound Infections Hayley Beth Schultz, Roshan B Vasani, Amy M. Holmes, Michael S. Roberts, and Nicolas H. Voelcker ACS Appl. Bio Mater., Just Accepted Manuscript • DOI: 10.1021/acsabm.8b00577 • Publication Date (Web): 07 Jan 2019 Downloaded from http://pubs.acs.org on January 13, 2019

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ACS Applied Bio Materials

Stimulus-Responsive Antibiotic Releasing Systems for the Treatment of Wound Infections Hayley B. Schultz a, b , Roshan B. Vasani a, c, Amy M. Holmes b, Michael S. Roberts b, d, e and Nicolas H. Voelcker*a, c, f, g a

Future Industries Institute, University of South Australia, Mawson Lakes, SA 5095, Australia b

School of Pharmacy and Medical Sciences, University of South Australia Cancer Research Institute, University of South Australia, Adelaide, SA 5001, Australia c

Monash Institute of Pharmaceutical Sciences, Monash University, Parkville, VIC 3052, Australia

d Diamantina

Institute, The University of Queensland, Translational Research Institute, QLD 4102, Australia

e School

of Pharmacy and Medical Sciences, University of South Australia, Basil Hetzel Institute

for Translational Medical Research, The Queen Elizabeth Hospital, Adelaide, SA, Australia f Melbourne

Centre for Nanofabrication, Victorian Node of the Australian National Fabrication Facility, 151 Wellington Road, Clayton, VIC 3168, Australia

g Commonwealth

Scientific and Industrial Research Organisation, Clayton, VIC 3168, Australia

*Corresponding Author. Telephone: +61 3 99039230. Email: [email protected]

ACS Paragon Plus Environment

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ACS Applied Bio Materials 1 2 3 4 5 6 7 8 9 10 11 12 13 14 15 16 17 18 19 20 21 22 23 24 25 26 27 28 29 30 31 32 33 34 35 36 37 38 39 40 41 42 43 44 45 46 47 48 49 50 51 52 53 54 55 56 57 58 59 60

ABSTRACT

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There remains an unmet need for innovative treatments for chronic wound

infections as they continue as a financial and social burden on society. Due to the dynamic nature of wounds, this study investigated the utilization of stimulus responsive plasma polymers for the development of pH- and thermo-responsive antibiotic delivery systems for the treatment of wound infections. Porous silicon films were loaded with the antibiotic levofloxacin (LVX), and subsequently coated with plasma polymer layers: first poly(1,7-octadiene) (pOCT) for stability, followed by either the temperature responsive polymer poly N,N-diethylacrylamide (pDEA) or the pH responsive polymer poly 2-(diethylamino)ethyl methacrylate (pDEAEMA), to fabricate two delivery systems. The delivery systems were thoroughly characterized chemically and physically, and tested in vitro through drug release and bacterial zone of inhibition studies. After a 16 hour time point, the system containing pDEA achieved 3.2-fold greater release at 45 °C compared to 22 °C, while the system containing pDEAEMA achieved 2.2-fold greater release when exposed to pH 8.5 media compared to pH 6.2 media. Furthermore, both systems retained their antimicrobial activity and demonstrated stimulus-responsive release to form zones of inhibition on relevant wound pathogens, Pseudomonas aeruginosa, Staphylococcus epidermidis and Staphylococcus aureus. Therefore, this proof-of-principle study confirms that stimulus-responsive porous silicon films can be utilized to deliver antibiotic when exposed to physiologically relevant stimuli such as pH and temperature with the potential to be applied to other pharmaceutics.

KEYWORDS

Levofloxacin; stimulus-responsive polymers; antibiotics; wound infections;

drug delivery systems; porous silicon.


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1. INTRODUCTION The cost of caring for and treating chronic wound patients in the US is estimated to be in excess of $25 billion per year.1 These include venous ulcers, diabetic foot ulcers and pressure ulcers and are most common in the elderly with either vascular disease or diabetes mellitus or who undergo surgical procedures.2 Wounds often become chronic upon infection, which impairs its ability to heal. The abundance and diversity of the pathogens that infect a wound is dependent on its type, depth, location and tissue perfusion.3 Often the infectious pathogens originate from the environment (air or from the cause of injury), the surrounding skin (the normal skin microflora), and endogenous sources involving mucous membranes (such as gastrointestinal or oral mucosae).34

Bowler et al. suggests that the primary causes of delayed healing and infection in both acute and

chronic wounds are due to aerobic pathogens such as Staphylococcus aureus, Pseudomonas aeruginosa, and beta-hemolytic streptococci.3 There is an unmet medical need for new treatments to cure wound infections. Current treatments can involve debridement of the infected tissue, use of dressings that contain therapeutics or applying topical antibiotics.5 It is vital that the infection is treated early to avoid the formation of antibiotic resistant bacterial biofilms, however it is difficult to determine whether a wound is infected or not in the early stages of infection.6 Additionally, misuse of antibiotics on uninfected wounds leads to strains of resistant microorganisms. Hence, this study investigates the development of a material to be used in wound dressings that would release antibiotics upon infection for early treatment and to ultimately cure the chronic wound infection. The wound environment is dynamic, and infection causes properties to change within the wound, including temperature and pH. Upon infection, the local temperature of the wound rises due to an increase in blood flow to the surrounding tissue.7 Additionally, the pH of an open wound can range


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from pH 6.5-8.88 and can become more acidic or more basic upon infection, depending on the type of bacteria that infects the wound bed.8-11 Therefore, wound pH and temperature are measures that can be exploited as indicators of infection.12-13 Controlled release drug delivery systems are designed to deliver the drug to the target site at a desired rate.14-15 While research has led to the development of wound dressings for direct placement on a wound for continual release of therapeutics to treat an infection,16-17 research investigating the controlled release of therapeutics from wound dressings in response to infection due to changes in the wound environment, i.e. temperature and pH, is scarce.6 Further research into an infection responsive antibiotic releasing material could lead to the effective early treatment of infection. Stimulus-responsive polymers change configuration in response to specific stimuli and have potential use in developing stimulus-responsive materials for treating wound infections. Poly N,Ndiethylacrylamide (pDEA) is a thermo-responsive polymer which exhibits a lower critical solution temperature (LCST) of 32-34 °C in an aqueous environment.18-19 An increase in temperature above the LCST leads to higher entropic contribution to the free energy and causes the polymer to collapse into a much more compact globule configuration.20-21 Plasma polymerization of pDEA has previously been investigated for the production of thermosensitive coatings on various substrates.22 Poly 2-(diethylamino)ethyl methacrylate (pDEAEMA) is a weak polybase with a pKa of 7.3.23 At a pH below the pKa, the amino group of the monomer units (displayed in Figure 1) within the polymer are protonated and repel adjacent monomers causing the polymer chain to expand.24 In contrast, at a pH above the pKa, the monomers are deprotonated, possess more intramolecular interactions causing the polymer to collapse.24 The plasma polymerization characteristics of pDEAEMA on glass slides has been previously described.25 Such stimulusresponsive polymers can be deposited directly onto surfaces. However, for the purpose of adhesion


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a non-stimulus responsive polymer such as poly(1,7-octadiene) (pOCT) is often deposited first as to stabilize the subsequent polymers and to sustain drug release of hydrophilic drugs.6, 26-27 As the LCST of pDEA and the pKa of pDEAEMA, respectively are similar to those of wound, they possess potential for use in stimulus-responsive drug delivery to infected wounds. Porous silicon (pSi) is a well-established nanostructured material that is both biocompatible and biodegradable.28 It is an optimal drug delivery scaffold as it possesses a high surface area for increased loading, and the physical parameters such as pore size and depth can be easily tuned to accommodate a wide range of therapeutics. Consequently, the material has been frequently studied as a drug carrier.2930

The fluoroquinolone antibacterial agent, levofloxacin (LVX) is a first line antibiotic, prescribed

systemically or topically to treat wound infections.31-32 It is sparingly soluble in water, has a pKa of 8.15 and is easily detected using fluorescence spectroscopy.33-34 The current study is a proof of concept investigation to explore the potential of the mentioned materials (stimulus-responsive and stabilizing polymers, LVX and pSi) in fabricating a stimulusresponsive system that can ultimately be incorporated into dressings to treat early wound infections. We exploit the pSi as a drug reservoir to load the antibiotic LVX, cap the drug loaded pSi layer with pOCT to control the release of the drug and apply a subsequent stimulus-responsive polymer layer (pDEAEMA or pDEA) to impart stimulus responsive properties to the drug releasing system. Two stimulus-responsive drug delivery systems were fabricated, one thermoand one pH-responsive, and their respective efficacies were tested against common wound pathogens.


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2. MATERIALS AND METHODS 2.1 Chemicals and Buffers. 2.1.1

Chemicals

Phosphate buffered saline (PBS) tablets, Trizma base, Trizma hydrochloride, imidazole, 1,7octadiene (OCT), acetone, 2-(diethylamino)ethyl methacrylate (DEAEMA), and levofloxacin (LVX) were purchased from Sigma-Aldrich (Castle Hill, Australia). Hydrofluoric acid solution 48% (HF) and hydrochloric acid (HCl) 37% were purchased from Scharlau (Port Adelaide, Australia). Ethanol and dimethylformamide (DMF) (RCI Labscan Limited) were obtained from Chem-Supply (Gillman, Australia). N,N-diethylacrylamide (DEA) was purchased from Polysciences Inc. (Gymea, Australia). Oxoid physiological saline solution 0.9%, Oxoid tryptone soya agar (TSA) and Oxoid tryptone soya broth (TSB) were purchased from Thermo Scientific (Thebarton, Australia). 2.1.2

Buffers

Milli-Q water (18 M cm; obtained from a Merck purification system) was used in the preparation of all buffers and in all experiments. The pH 7.4 PBS buffer was prepared as per the instructions on the bottle. Trizma base and Trizma hydrochloride were dissolved in water to yield a 0.2 M pH 8.5 buffer at 37 C. 0.2 M imidazole and 0.2 M HCl were combined with water to yield a pH 6.2 buffer at 25 C. Herein, the pH 7.4 PBS buffer, the pH 8.5 Trizma buffer and the pH 6.2 imidazole/HCl buffer will be referred to as the pH 7.4 buffer, the pH 8.5 buffer and the pH 6.2 buffer, respectively. 2.2 Sample Production


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The overall fabrication process is summarized in Figure 1, illustrating the final products, LVX loaded pSi coated with pOCT and pDEAEMA (pSi-pOCT-pDEAEMA) and LVX loaded pSi coated with pOCT and pDEA (pSi-pOCT-pDEA).


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Figure 1. The stimulus-responsive drug delivery system fabrication process. (A) A silicon wafer is electrochemically etched to produce pSi and is oxidized. (B) pSi is loaded with LVX by means of spin coating. (C) Plasma polymer deposition is used to deposit pOCT on the surface. (D) Plasma polymer deposition is used to deposit the pH-responsive pDEAEMA or (E) the thermo-responsive pDEA onto the surface. The chemical structures of the monomers are included. 2.2.1

Electrochemical Etching and Oxidation of Silicon

P-type silicon wafers, sourced from Siegert Wafer (Aachen, Germany), were used for etching. The wafers were boron doped, had an orientation of ±0.5°, a resistivity of 20 mg/mL LVX in 0.9% saline solution, and negative control, 0.9% saline solution, were also used to ensure the validity of the experiment. 10 µL aliquots of the standards and controls were applied to the paper discs on the inoculated agar plates, inverted and incubated at 37 C for 18 h. The range of concentrations were expected to cover the range of expected minimum inhibitory concentration (MICs) for the three bacterial strains. 37-38 The MIC range for each bacterial strain was determined as being between the lowest concentrations of LVX that exhibited a zone of bacterial growth inhibition (ZOI) and the highest concentrations of LVX that exhibited no ZOI.


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2.6.3

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Antimicrobial Action of LVX Eluted from the Systems

LVX was extracted from 0.5 cm2 pieces of the systems by incubating the samples in 110 µL of their relevant buffers and temperatures (see Section 2.4.1) for 1, 3 and 16 h. Extractions were taken after 1 and 3 h to confirm that the release of the antibiotic and its effects occurred quickly after applying the dressing, and after 16 h. This time point was chosen since dressings are typically changed 1-2 times a day. 10 µL aliquots of the eluted LVX solutions were applied to the paper discs on the inoculated agar plates, which were then inverted and incubated at 37 C for 18 h. A positive control, >20 µg/mL LVX in 0.9% saline solution, and negative controls, pH 6.2 buffer, pH 8.5 buffer and pH 7.4 buffer, were also applied to discs to ensure the validity of the experiment. Subsequent to incubation, the ZOI were calculated by measuring the diameter (mm) of the ZOI and subtracting 5 mm to compensate for the 5 mm paper disc being used.


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3. RESULTS AND DISCUSSION In this proof-of-concept study, two stimulus-responsive drug delivery systems were fabricated, pH-responsive pSi-pOCT-DEAEMA and thermo-responsive pSi-pOCT-DEA. These systems were developed for the potential application of incorporation into wound dressings to impart infection triggered antibiotic release properties to treat wound infections. Here, we will describe the physicochemical characteristics of the systems, demonstrate their stimulus-responsive properties, compare their triggered release and finally, investigate the bioavailability and antimicrobial effectiveness of the released LVX on common wound pathogens. 3.1 Physicochemical Characterization Silicon wafers were electrochemically etched to produce pSi films. Pieces of pSi films were subsequently loaded with levofloxacin using the spin coating method, followed by the deposition of the plasma polymers pOCT and pDEAEMA or pDEA. After fabrication of the two stimulusresponsive antibiotic releasing materials, pH-responsive pSi-pOCT-pDEAEMA and thermoresponsive pSi-pOCT-pDEA, the materials were thoroughly characterized both physically and chemically. Physical characterization included SEM and AFM and chemical characterization included ATR FT-IR and XPS, and were used to confirm the successful deposition of plasma polymer layers. 3.1.1

Scanning Electron Microscopy (SEM)


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The SEM images are displayed in Figure S1. The images confirmed the width of pore openings of the pSi to be approximately 11.08 ± 0.90 nm (mean ± SD, n=3) and the depth of the etched layer to be approximately 5.86 ± 0.01 m (mean ± SD, n=3). Plasma polymerization resulted in full coverage of the pSi surface with little penetration into the porous layer. Distinct pOCT, PDEAEMA and pDEA layers were observed (see Supplementary Information for more details). 3.1.2

Atomic Force Microscopy (AFM)

The deposition rates of the polymers were determined by using AFM. The polymers were deposited on glass cover slips for a known period of time, carefully scratched with a scalpel and then imaged using AFM. The measured thicknesses divided by the deposition times yielded the deposition rates. pOCT, pDEAEMA and pDEA were deposited at rates of 7.55, 2.13 and 0.96 nm/min, respectively. It has been reported that deposition rates of plasma-polymerized polymer films are proportional to the polymerization time and have a linear relationship.39 Therefore, these rates were used to calculate the polymerization time required to obtain the desired polymer thickness as shown in Table S1. 3.1.3

Attenuated Total Reflectance Fourier Transform Infra-red (ATR FT-IR)

The surfaces of the delivery systems were characterized with ATR FT-IR after each plasma polymer deposition to confirm the successful functionalization (Figure 2). The peaks at approximately 1030 cm-1 and 800 cm-1 were present in all four spectra, and correspond to Si-O-Si and –O-Si-H vibrations from the oxidized pSi layer, respectively.40-42 While the spectrum of the oxidized pSi (spectrum A) contained Si-O peaks, it did not contain any other peaks of significance.43 Subsequent to the deposition of pOCT, four new peaks were visible in the pSi-pOCT spectrum (spectrum B). These new peaks at 2951 cm-1 and 2925 cm-1 corresponded to asymmetric and


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symmetric C-H stretching44 and 1462 cm-1 and 1384 cm-1 corresponded to C-C stretching vibrations of the –C-CH3 and -CH2- groups.

44-46

The same peaks were also visible after the

deposition of pDEAEMA (spectrum C) and pDEA (spectrum D). pOCT monomer was also expected to contain C=C bonds with a peak at 1609 cm−1. However, the absence of this peak in the measured spectra indicates complete polymerization of the monomer through the double bond in the deposited film.22 In addition to the peaks already mentioned, the spectra after the addition of pDEAEMA (spectrum C) and pDEA (spectrum D) also contained other peaks. The spectrum of DEAEMA is known to possess a peak at 1730 cm-1 due to the C=O of the ester functional group. 44-46 This peak is seen in spectrum C, however an additional peak at 1643 cm-1 peak indicates the presence of C=O from a carboxylic acid or aldehyde group, likely due to fragmentation of the DEAEMA monomer through the ester group during plasma deposition.22 Additionally, pDEAEMA surfaces are known to display a peak at approximately 1155-1181 cm-1, corresponding to the C-N of the amine.44 This region was within the large Si-O peak region of the spectrum and therefore it is expected that it is hidden by the Si-O peak. Spectra D contained a peak at 1643 cm-1 which was assigned to the DEA amide.22 Finally, the broad –OH peaks, at 3307 and 3365 cm-1 on spectra C and D, respectively, are common for all polymer samples that have been exposed to air. However, the peak was absent from spectrum B and is thought to be due to the hydrophobicity of the pOCT layer. The presence of these peaks analyzed by ATR FT-IR confirmed the successful deposition of pOCT, pDEA and pDEAEMA on pSi.


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Figure 2. ATR FT-IR spectra of (A) pSi, (B) pSi-pOCT, (C) pSi-pOCT-pDEAEMA and (D) pSipOCT-pDEA. 3.1.4

X-Ray Photoelectron Spectroscopy (XPS)

XPS also confirmed the presence of the plasma polymer layers. The elemental compositions of the surfaces derived from the XPS survey spectra are displayed in Table 1. As expected from the chemical formulae for the OCT (C8H14), DEAEMA (C10H19NO2) and DEA (C7H13NO) monomers, carbon peaks (285 eV) were observed in all spectra, and additional oxygen (531 eV) and nitrogen peaks (402 eV) were observed in the spectra of pDEAEMA and pDEA which is consistent with literature.21 Oxygen (531 eV) and Si peaks (105 and 156 eV) were observed in the oxidized pSi spectra 47 and the measured surface composition was in good agreement with the literature.48 The presence of carbon on this surface indicates that some atmospheric contamination had taken place. Following the deposition of pOCT, the carbon content increased to 95.9% due to the hydrocarbon


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nature of pOCT. However, 4.1% oxygen was present, likely due to the reaction between atmospheric oxygen and long-lived reactive species within the polymer surface, which has been described by Whittle et al. (2000) who measured up to 5% oxygen content.35 Following the deposition of pDEAEMA on the pSi-pOCT surface, the measured surface composition was somewhat different to the theoretical surface composition. In the literature, it has been reported that XPS analysis of pDEAEMA has resulted in surface composition of 74.9% carbon, 7.7% nitrogen and 17.4% oxygen.49 The measured value of carbon, 83%, was greater than the theoretical and literature values,49 while the nitrogen and oxygen percentages were less than expected, 4.8% and 11.2%, respectively, perhaps due to fragmentation caused by plasma polymerization22, 50 or due to the influence of the underlying pOCT layer. The measured surface composition for pSipOCT-pDEA was is in relatively close agreement with the theoretical surface composition and literature values.22 No silicon signals were present in the spectra of the polymerized samples, suggesting complete surface coverage of the pSi by the polymer films and polymer layers thicker than 10 nm (the approximate depth of the XPS analysis).21-22

Table 1. Theoretical and measured surface compositions from XPS measurements. A dash (-) indicates no theoretical concentration or no measured concentration. Surface Concentrations (atomic %) Surface

Theoretical

Measured

C

N

O

Si

C

N

O

Si

pSi-pOCT

100.0

-

-

-

95.9

-

4.1

-

pSi-pOCTpDEAEMA

76.9

7.7

15.4

-

83.0

4.8

11.2

-

pSi-pOCT-pDEA

77.8

11.1

11.1

-

76.0

14.5

9.5

-

Oxidized pSi

-

-

-

-

4.4

-

63.5

32.1


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High-resolution C 1s XPS scans were also collected to examine the chemical bonding states of carbon in each of the polymer films to confirm their deposition and to supplement the ATR FT-IR data. The high-resolution C 1s XPS spectra together with spectral deconvolutions of the systems are displayed in Figure S2 and were consistent with the literature (see supplementary information for more details). This collection of XPS results, together with the results from ATR FT-IR, provides strong evidence that all polymers were successfully deposited onto the surface of pSi. 3.2 Stimulus-Responsive Behavior of the Material Surfaces After confirming the deposition of the plasma polymers on pSi, the stimulus-responsive behavior of the surfaces were studied. This was accomplished by measuring the difference in CA’s and IRS EOT ‘switches’ to demonstrate the pH- and thermo-responsive natures of pSi-pOCT-pDEAEMA and pSi-pOCT-pDEA, respectively. 3.2.1

Static Water Contact Angle (CA) Measurements

3.2.1.1 Change in CA Values with each Functionalization Step The CA values between droplets of buffers and the surface of systems were measured after each polymerization step, to confirm the success of the plasma polymer deposition (Figures 3.A and 3.B). For the pSi-pOCT-pDEAEMA and the pSi-pOCT-pDEA systems, pH 6.2 and pH 7.4 buffers were used, respectively, at RT. These conditions represent the systems without the presence of the stimuli. Figure 3.A presents the CA’s between the DEAEMA sample surfaces and pH 6.2 buffer at RT. The oxidized pSi displayed the smallest contact angle, 17°, and therefore was the most hydrophilic of the surfaces. This was due to the highly hydrophilic SiO2 surface strongly interacting with the buffer causing it to spread out on the surface.51 This result was in close agreement with the


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literature value of 20 using water.51 After the deposition of pOCT, the pSi-pOCT surface had the largest CA, 71°, and was therefore the least hydrophilic of the surfaces. This is due to the hydrophobic nature of the hydrocarbon rich pOCT polymer layer.52 Literature states a range of CA values (80 - 92) for pOCT.53-56 The measured value lies below this range, however this may be explained by the water penetrating the hydrophobic pOCT layer (located on top of the pores) into the very hydrophilic oxidized pores of the Si leading to better wetting. The Cassie-Baxter effect exhibited by the underlying pSi layer should also be taken into consideration. The Cassie-Baxter effect generates higher apparent CA values on rough surfaces due to heterogeneous wetting caused by pockets of vapor caught underneath the droplet of liquid producing a composite interface where less area is in contact with the solid substrate.57 It is expected that the Cassie-Baxter effect is generating an effect here, though the net result of its effect in combination with the hydrophilic porous under layer is likely leading to the lower contact angles observed. Following the deposition of pOCT, a more hydrophilic layer of pDEAEMA was deposited, decreasing the contact angle to 47. In comparison to pSi-pDEAEMA, pSi-pOCT-pDEAEMA had a CA that was 3° greater due to the presence of the underlying hydrophobic pOCT layer. The same trends were seen with the CA’s between the DEA samples and pH 7.4 buffer at RT in Figure 3.B. The oxidized pSi had the smallest CA, 18°, pSi-pOCT had the largest contact angle, 66°, and the pSi-pOCT-pDEA surface was less hydrophilic than the oxidized pSi and more hydrophilic than pOCT with a contact angle of 39°. pSi-pOCT-pDEA had a CA that was 2° greater than pSi-pDEA due to the presence of the underlying hydrophobic pOCT layer. Finally, the 5° greater CA seen with pH 6.2 buffer on pSi-pOCT compared to pH 7.4 buffer on pSi-pOCT may be contributable to the difference in buffer pH and composition. 3.2.1.2 Change in CA Values with Stimulus


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Additionally, CA values were measured in the presence of the relevant stimulus to show how the stimulus generates a change in CA (Figures 3.C and 3.D). The CA values between pH 6.2 buffer and pH 8.5 buffer were compared on the surface of the pSi-pOCT-pDEAEMA, whereas the CA values between pH 7.4 buffer at RT and 45° were compared on the surface of the pSi-pOCTpDEA. In Figure 3.C, the pOCT did not show pH-responsive nature. When the pH of the buffer changed from 6.2 to 8.5, only a very small change in contact angle was observed of 1 and was not statistically significant (p = 0.5, n=3). In contrast, the pSi-pOCT-pDEAEMA did show pHresponsive nature as when the pH changed, there was a clear increase in contact angle of 5 with statistical significance (p

Stimulus-Responsive Antibiotic Releasing Systems for the Treatment

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