Double-Walled Microparticles-Embedded Self-Cross-Linked, Injectable, and Antibacterial Hydrogel for Controlled and Sustained Release of Chemotherapeutic Agents.

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Double-walled Microparticles-embedded Self-crosslinked, Injectable, and Anti-bacterial Hydrogel for Controlled Sustained Release of Chemotherapeutic Agents Pooya Davoodi, Wei Cheng Ng, Wei-Cheng Yan, Madapusi P. Srinivasan, and Chi-Hwa Wang ACS Appl. Mater. Interfaces, Just Accepted Manuscript • DOI: 10.1021/acsami.6b03041 • Publication Date (Web): 17 Aug 2016 Downloaded from http://pubs.acs.org on August 17, 2016

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ACS Applied Materials & Interfaces

Double-walled Microparticles-embedded Self-crosslinked, Injectable, and Anti-bacterial Hydrogel for Controlled and Sustained Release of Chemotherapeutic Agents

Pooya Davoodi 1, Wei Cheng Ng 2, Wei Cheng Yan 1, Madapusi P. Srinivasan 1,3, Chi-Hwa Wang 1*

1

Department of Chemical and Biomolecular Engineering, National University of Singapore, 4

Engineering Drive 4, Singapore 117585. 2

NUS Environmental Research Institute, National University of Singapore, 1 Create Way, Create Tower

#15-02, Singapore 138602. 3

Civil, Environmental and Chemical Engineering, RMIT University, GPO Box 2476, Melbourne VIC

3001, Australia.

*

Corresponding author: Tel.: +65-65165079; Fax: +65-67791936 (C.H. Wang),

E-mail address: [email protected]

1 ACS Paragon Plus Environment


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ABSTRACT First-line cancer chemotherapy has been prescribed for patients suffered from cancers for many years. However, conventional chemotherapy provides a high parenteral dosage of anticancer drugs over a short period, which may cause serious toxicities and detrimental side effects in healthy tissues. This study aims to develop a new drug delivery system (DDS) composed of double-walled microparticles and an injectable hydrogel for localized dual-agent drug delivery to tumors. The uniform double-walled microparticles loaded with cisplatin (Cis-DDP) and paclitaxel (PTX) were fabricated via coaxial electrohydrodynamic atomization (CEHDA) technique and subsequently were embedded into injectable alginate-branched polyethyleneimine. The findings show the uniqueness of CEHDA technique for simply swapping the place of drugs to achieve a parallel or a sequential release profile. This study also presents the simulation of CEHDA technique using computational fluid dynamics (CFD) that will help in the optimization of CEHDA’s operating conditions prior to large-scale production of microparticles. The new synthetic hydrogel provides an additional diffusion barrier against Cis-DDP and confines premature release of drugs. In addition, the hydrogel can provide a versatile tool for retaining particles in the tumor resected cavity during the injection after debulking surgery and preventing surgical site infection due to its inherent antibacterial properties. Three-dimensional MDA-MB231 (breast cancer) spheroid studies demonstrate a superior efficacy and a greater reduction in spheroid growth for drugs released from the proposed composite formulation over a prolonged period, as compared with free drug treatment. Overall, the new core-shell microparticles embedded into injectable hydrogel can serve as a flexible controlled release platform for modulating the release profiles of anti-cancer drugs and subsequently providing a superior anticancer response.


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Keywords: Injectable hydrogels; Core-shell microparticles; Antibacterial surfaces; Dual-agent drug delivery; 3D tumor spheroids

1. INTRODUCTION Hydrogels are an important class of materials with unique features for biomedical applications exhibiting an exceptional biocompatibility given their physiochemical properties and high capacity for water absorption.

1-3

In addition, hydrogels can provide a 3-dimentional

microstructure, which closely mimics natural extracellular matrix of cells for tissue regeneration. 4-6

The mechanical and physiochemical characteristics of hydrogels can be readily modified for

clinical practices. For instance, the permeability of hydrogels can be readily manipulated through the control of pore size and degradation rate, where an optimal porous structure benefits the transport of medications and nutrients inside the gel network.

7

Leveraging these fascinating

properties, researchers have developed numerous hydrogel systems for various clinical applications ranged from drug delivery and tissue engineering to diagnostics and coating of medical implants. 8 Recently, injectable in situ-gelling hydrogels have gained more attention due their superior advantages compared with the implantable counterparts.

9

The macroscopic dimension of

preformed hydrogels is the main obstacle in the use of hydrogel implants in a non-invasive surgery. However, injectable hydrogels can facilitate the administration and protection of bioactive molecules or cells via simple and non-invasive approaches.10-12 The use of injectable hydrogels not only enhances patient compliance but also reduces recovery time, risk of infection 13-14

, and treatment costs.

15

Although the highly porous structure and hydrophilic nature of

hydrogels benefit their widespread applications for tissue engineering, premature release of



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entrapped molecules (e.g. anti-cancer drugs) from hydrogel confines their utilization as drug delivery vectors. Bulk modification of hydrogels (e.g. increasing cross-link density) can partially diminish this limitation, while it may have adverse effects on biocompatibility and mechanical properties of the hydrogel.

16-18

On the other hand, a composite hydrogel system with the

immobilized hydrophobic depots inside a hydrophilic gel matrix may overcome the physiochemical limitations of hydrogels for the encapsulation and controlled release of small molecules. 19-20 Over the past two decades, polymeric microparticles loaded with different medications for local drug delivery have demonstrated versatile features in the treatment of chronic disease (e.g. cancer).

21-24

Microparticles can locally release their payload at tumor site, which significantly

reduces detrimental systemic side-effects and improves patients’ satisfaction and compliance. However, the use of single agent therapy often fails to achieve complete cancer remission due to rapid development of drug resistance in tumor cells. Recently, combination chemotherapy has opened a new avenue for the treatment of drug-resistant tumor cells.

25-26

The use of multiple

drugs with different mechanisms of action has exhibited less cancer adaptation process (i.e. development of drug resistance through abnormal altered drug transport) and remarkable therapeutic outcome through synergistic effects against cancer cells. Therefore, in addition to encapsulating chemotherapeutic agents, microparticles should be engineered for the controlled release of multiple agents at target tissues with optimal release rates. 27-29 Microparticle-based local drug delivery is a promising technique to reduce the risk of recurrence for cancer survivors after a segmental resection. Drug loaded microparticles can provide a high local concentration gradient of drugs within 2 cm of a resected tumor and target tumor cells remained after a surgery. However, retaining the particles (in suspension) in the


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surgical cavity during the injection and treatment period is still a major challenge confining the wide use of such a delivery system.

19-20

Therefore, a 3D biodegradable matrix, which can

efficiently entrap and retain microparticles in the cavity is strongly needed regulating of release profiles

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, participate in

, and fill the cavity after tumor resection. An in situ-gelling

hydrogel would be the best candidate that can play all the aforementioned roles successfully. The pre-encapsulation of therapeutic medications into micro-particulate structures and the subsequent entrapment of the depots inside a hydrogel system can elevate drug loading and delivery duration in a controlled manner, while immobilizing delivery vectors at a target site. We have reported that double-walled microparticles consisting of a polymeric core surrounded by a shell layer are able to concurrently encapsulate hydrophilic and hydrophobic drugs and release them over prolonged period in a controlled manner. 32-33 In this study, a novel drug delivery system composed of microparticles and an injectable hydrogel is introduced to address the shortcomings mentioned above. We employed coaxial electrohydrodynamic atomization (CEHDA) for the fabrication of double-walled microparticles loaded with Cis-DDP and PTX in the core and shell compartments, respectively. Particles were examined for drug loading, encapsulation efficiency, and their ability for controlled release of the therapeutic agents. The interior geometry of the particles was also examined using laser scanning confocal microscopy. Next, these particles were entrapped inside an in situ-gelling carbohydrate-based hydrogel to form a micro-composite structure. The hydrogel physiochemical properties such as degree of swelling and degradation rate, antibacterial effects, cellular cytotoxicity, and its impact on the release kinetics of both agents were investigated and compared with the drug release kinetics achieved by microparticles alone.



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2. MATERIALS AND METHODS 2.1. Materials Poly(lactic-co-glycolic acid) (PLGA 50:50, inherent viscosity in hexafluoroisopropanol 0.550.75 dL g-1) and poly(D,L-lactide) (PDLLA, inherent viscosity in hexafluoroisopropanol 0.55-0.75 dL g-1) were obtained from LACTEL Absorbable Polymers (DURECT Corporation, USA). Sodium alginate (medium viscosity of 3500 cps for a 2% in H2O), branched polyethyleneimine (25 kDa), sodium metaperiodate (NaIO4, ≥99.0 %), and picrylsulfonic acid (5% in H2O) were purchased from Sigma-Aldrich (Singapore). Paclitaxel was supplied courtesy of Bristol-Meyers Squibb (New Brunswick, NJ) and cis-diamminedichloroplatinum (II) (Cis-DDP, ≥99.99 %) was obtained from Sigma-Aldrich (Singapore). HPLC-grade dichloromethane (DCM) (TEDIA, Fairfield, OH, USA) and phosphate buffer saline (10×PBS) were used for the fabrication of microparticles and in vitro release tests, respectively. Dulbecco’s modified eagle medium (DMEM), fetal bovine serum (FBS), penicillin-streptomycin, and trypsin-EDTA (0.25%) were purchased from Life Technology (Singapore). All other materials and reagents used were of analytical grade and used without further purification.

2.2. Fabrication of double-layered microparticles The fabrication of double-layered microparticles comprising of PDLLA (core) and PLGA (shell) compartments, and loaded with drugs, were accomplished using CEHDA technique, as reported earlier.

32

Briefly, PDLLA and PLGA (8 wt.%, in DCM) solutions were prepared

overnight and mixed properly with Cis-DDP drug solution, where Cis-DDP was dissolved in water or dimethyl sulfoxide (DMSO) or dimethylformamide (DMF). PTX was directly dissolved in polymer solution at a pre-determined weight ratio. The polymer/drug solutions were


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subsequently pumped through a co-centric metal nozzle at predetermined flow rates for the core (0.5 ml h-1) and shell (2.5 ml h-1) streams, respectively. The solutions formed a liquid cone with a thin emerging jet at the tip of the nozzle due to increase of potential difference between the nozzle and grounded substrate using a high voltage generator (Glassman High Voltage Inc., NJ, USA). The polymer jet break-up occurred owing to accumulation of positive charges at liquid-air interface and formed mono-dispersed micro-droplets. Solid polymeric particles were collected on the negatively charged collectors after quick evaporation of DCM from droplets. Lastly, microparticles were freeze-dried for two days to fully eliminate residual DCM from samples and stored at ˗20°C for characterization and in-vitro experiments. To have a better understanding of the phenomena that occurred during particle fabrication, computational fluid dynamic (CFD) study was performed in the area near the nozzle tip. The equations used in CFD simulation were introduced in supplementary information (Table S1).

2.3. Aldehyde functionalized sodium alginate The aldehyde functionalization of sodium alginate was performed as previously described elsewhere.

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Briefly, sodium alginate (5.0 g) was dispersed in 25 ml of ethanol. Sodium

metaperiodate (2.5 g) was dissolved in 25 ml of ultra-pure water and was added dropwise to the suspension of alginate/ethanol. The reaction mixture was then stirred magnetically for 6, 12, and 18 h at room temperature in dark. The solution was then transferred into tubular dialysis membrane (MWCO= 1000 Da, Spectrum® Laboratories, INC), where it was extensively dialyzed against deionized-water (5 L). A minimum of five dialysis cycles were performed prior to lyophilization of the solution.



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2.4. Synthesis of in-situ gelling hydrogels and the micro-composite To prepare Alg-Ald (3 w/v %) and bPEI (1 %, 2.5%, 5%, and 10 w/v %) solutions, the chemicals were separately dissolved in ultra-pure water and the pH of the solutions was adjusted at ~ 7.5 using hydrochloric acid (HCl, 32.5%, VWR International). Hydrogels were formed by mixing an equal volume (150 µl) of Alg-Ald and bPEI solutions. The hydrogel was allowed to reach to stable form for 30 min at 37°C. For microparticles loaded hydrogels, microparticles at the final 25%, 50%, and 75% of dried-gel weight were suspended in Alg-Ald solution and quickly mixed with bPEI solution to form hydrogel composites. The hydrogel is formed upon mixing bPEI (1%) and Alg-Ald (3%) solutions (short gelation time, ~2-3 seconds). However, the gelation time becomes longer as bPEI concentration increases, where mixing bPEI (10%) and Alg-Ald (3%) solutions does not produce obvious hydrogel pellet (the viscosity of the mixture increases, but no obvious hydrogel is formed).

2.5. Physicochemical characterization 2.5.1. Morphology, particle size, and uniformity The surface morphology and size distribution of microparticles were examined via Scanning Electron Microscopy (SEM) (JEOL JSM 5600LV, Tokyo, Japan) with an accelerating voltage of 10 kV. The microparticles were collected on the surface of a double-side adhesive carbon tape and subsequently coated with a thin layer of platinum using a JFC-1300 Platinum Coater (JEOL, Tokyo, Japan) at 30 mA for 50 sec prior to analysis. The inner morphology of microparticles was observed using Confocal Laser Scanning Microscopy (CLSM) (Olympus FluoView™ FV1000, Olympus America Inc., USA), where


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coumarin 6 (Sigma-Aldrich, Singapore), a green fluorescence dye, was mixed with the polymer solution of the inner nozzle prior to fabrication of microparticles.

2.5.2. Drug loading and encapsulation efficiency The drug loading and encapsulation efficiency were determined by dissolving 15 mg of microparticles in 2 ml DCM and were allowed to stand until complete disappearance of the microparticles. Then, 5 ml 1×PBS was added to the previous solution and the mixture was incubated at 37°C overnight. The concentrations of Cis-DDP and PTX were determined using inductively coupled plasma optical emission spectrometry (ICP-OES) and high-performance liquid chromatography (HPLC), respectively. The HPLC apparatus was equipped with a C18 reverse-phase (RP) column and UV-2487 UV-detector where acetonitrile: H2PO4 buffer (50:50 v/v) at 1.0 ml min-1 was utilized as mobile phase.

2.5.3. Determination of the degree of crosslinking The cross-linking degree of the hydrogel (Ald-Alg: bPEI of 3:1 w/w %) at different pH was determined using picrylsulfonic acid (2,4,6-trinitrobenzene sulfonic acid (TNBSA)) reagent, known as a rapid and sensitive technique for quantitating unreacted primary amines. After the formation of hydrogels, the samples were centrifuged and 10 µl of each supernatant was diluted with 90 µl of 0.1M sodium bicarbonate (pH~8.5) before being transferred (25 µl) into 96-well plate. Next, 50 µl of freshly prepared TNBSA (0.01 w/v %) reagent was introduced to the diluted samples and incubated at 37°C for 2 h, required for the completion of the reaction. The reaction was terminated after adding 25 µl of 10% sodium dodecyl sulfate (SDS) and 12.5 µl of 1N HCl to each sample. The absorbance of the mixtures was measured at a wavelength of 335 nm using



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Infinite® 200 PRO microplate reader (TeCAN, Switzerland). The amount of unreacted bPEI was calculated after comparing the results with a standard curve of serially diluted pure bPEI solutions.

2.5.4. Fourier transform infrared spectroscopy (FTIR) and solid-state 13C-NMR Fourier transform infrared spectrum of Alg-Alds and final hydrogel products were recorded using Bio-Rad FTS-3500ARX at transmittance mode in the range of 400 cm-1 to 4000 cm-1. Potassium bromide (KBr) pellets were used for the preparation of samples. A bare KBr pellet was considered as background, where its corresponding spectrum was subtracted from that of samples. All samples were purged by nitrogen atmosphere for 15 min before recording their absorbance. The chemical structure of the hydrogel and the formation of Schiff’s base chemical bonds were determined using solid-state

13

C-nuclear magnetic resonance (13C-NMR)

spectroscopy (BRUKER 400 MHz, equipped with a 4 mm MAS BB-BB probe) at room temperature (296.1 K). The frequency and the number of scans were set at 100 MHz and 16380, respectively.

2.5.5. Gel permeation chromatography (GPC) Relative molecular weight and weight distribution of the native and oxidized alginates were measured by gel permeation chromatography system equipped with a PL aquagel-OH Mixed 300 column (300×7.5mm, 8u) and Waters 2414 Reflective Index detector at constant temperature (35 º

C). Water was used as mobile phase at a flow rate of 1.0 ml min-1. Sample injection was set at

20 µl. A set of five polyethylene glycol (PEG) solution with standard molecular weights was used for molecular weight calculation.


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2.5.6. Elemental analysis The percentage estimation of three key elements (i.e. carbon (C), hydrogen (H), and nitrogen (N)) in the hydrogel samples were measured using Elementar vario MICRO cube analyzer. All samples were freeze-dried for 48 h prior to analysis.

2.5.7. Thermogravimetric analysis (TGA) TGA was performed to investigate the thermal profile of intermediates and final hydrogel products, where bare sodium alginate was used as control sample. Samples of 5-15 mg were prepared and freeze-dried overnight and subsequently loaded in thermogravimetric analyzer (DTG-60 AH, Simultaneous DTA-TG apparatus, Shimadzu) furnace equipped with TA-60WS thermal analyzer device. The samples were heated up to 600ºC under dry nitrogen gas at a flow rate of 10 ml min-1 and the measurements were conducted at the scanning speed of 10°C min-1. The weight loss pattern of each sample was recorded simultaneously during the analysis.

2.5.8. Differential scanning calorimetry The thermal properties of intermediates Alg-Ald and final hydrogel products and the physical state of the drug inside microparticles were determined using METTLER TOLEDO 822e (Greifensee, Switzerland) equipped with a gas controller (TSO800GC1) and controlled by STARe Default DB V9.10-STARe software. Approximately 5-10 mg of the microparticles were sealed in covered aluminum pans. A standard empty pan was used as a reference for all analyses. Samples were tested over the temperature range from ˗10°C to 110°C at a heating rate of 10 °C min-1. The equipment was purged with nitrogen gas at a rate of 50 ml.min-1 and a liquid nitrogen



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stream was employed as coolant. Blank microparticles were used as control in all thermal analysis tests.

2.5.9. Determination of swelling ratio The samples were prepared as described above and were freeze-dried overnight. A known weight of dry gels was soaked in ultra-pure water at room temperature. At pre-determined intervals, the excessive surface water was removed using filter paper. After weighting the samples, the swelling ratio was calculated using the following equation: SR (%) = (Wi,st /Wi,s0)×100

(1)

where Wi,st and Wi,s0 were the weight of sample i at time t and the dry sample, respectively.

2.5.10. Bulk degradation of hydrogels The hydrogels were freshly prepared and placed in Eppendorf test-tubes filled with 1.5 ml 1×PBS solution. The samples were incubated at 37 °C and 240 rpm and the degradation rates were determined at predetermined time intervals. The samples were lyophilized at indicated time-points and accurately weighted to calculate the weight loss for each sample.

2.5.11. Rheological measurement The rheological properties of bPEI (1%)/Alg-Ald (3%) hydrogel was investigated using ARG2 Stress Controlled Rheometer (TA Instruments, U.S.A.). Stress sweep and frequency sweep measurements were carried out at 37 °C using a parallel plate configuration (diameter, 20 mm) with a gap of 500 µm. The storage (elastic) modulus (Gʹ) and the loss of modulus (Gʺ) of the


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hydrogel were measured as functions of oscillatory pressure (0.5˗1000 Pa, 1.0 Hz) and angular frequency (~0.6˗600 rad s˗1, 0.1 kPa). All samples were tested ~1 h after preparation.

2.6. Bacteria culture The antibacterial property of the synthetic hydrogels was examined against Escherichia coli (E.coli, Gram negative) and Bacillus pumilus (B. pumilus, Gram positive) bacteria. Prior to the experiments, the bacteria were cultivated in a streaked agar-plate to form distinct colonies overnight. A single colony was harvested and subsequently inoculated in autoclaved tryptic soy broth (TSB). The bacteria were allowed to grow overnight at 37°C, shaking at 120 rpm to ensure that they grew at the exponential growth-phase. The bacteria concentration was monitored photometrically 600 nm using NanoDrop 2000 spectrophotometer (Thermo Fisher Scientific Inc., U.S.A). The initial optical density (O.D.) of the bacteria solution was adjusted to ~0.1, corresponded to concentrations ~108 CFU ml-1.

2.6.1. Concentration-dependent antibacterial activity Hydrogels used for the antibacterial experiments were freshly prepared by mixing an equal volume of bPEI solutions (100 µl) at different concentration with 100 µl of Alg-Ald (3 wt %, MMw, 6 h oxidation) solution. The hydrogels were then transferred to a 96-well plate and incubated at 37°C for 1 h before washing gently with sterilized 1×PBS. Next, the stock bacterial samples were further diluted to 104 CFU ml-1 in 1×PBS, followed by introducing 100 µl to the surface of the gels. Bacteria samples without hydrogels were employed as control. The samples were then incubated at 37 °C with constant shaking at 120 rpm for 18 h, after which 10 µl of the



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supernatant was serially diluted. The number of colony-forming units (CFUs) (on agar-plate) was determined and compared with control after incubation at 37°C for 12 h. To investigate the effect of bacteria concentration on the antibacterial behavior of the hydrogel, sample #1 (S1) was incubated with a 90 µl of bacteria solution prepared after 1:10 serial dilutions of a concentrated stock of bacteria (108 CFU ml-1) with TSB. The O.D. and the number of colony-forming units (CFUs) (on agar-plate) were determined and compared with control as previously explained.

2.6.2. Time-dependent antibacterial activity The killing kinetic of the hydrogel was determined via monitoring the concentration of bacteria solutions at indicated intervals. Briefly, the freshly prepared hydrogels were incubated with 100 µl of bacteria solution (104 CFU ml-1) at 37 °C, shaking at 120 rpm. At the specific time-points, 10 µl of the supernatant was removed, serially diluted in 1×PBS, and placed on an agar plate surface at 37 °C. The number CFUs was counted after 12 h of incubation.

2.6.3. Assessment of bacteria-hydrogel contact on antibacterial behavior Fresh hydrogels (400 µl of bPEI 1.0 wt. % +400 µl of Alg-Ald 3 wt. %) were washed with 1×PBS once, and transferred into transwell cell culture (Corning® Transwell®) inserted in a 24well plate. A volume of 400 µl of 104 CFU ml-1 bacteria (E.coli) solution in TSB was loaded to each well and an additional 100 µl of bacteria-free TSB was introduced on top of the gels in the transwell inserts. A transwell insert and bacteria solution without hydrogel were employed as negative controls. All samples were incubated at 37 °C with continuous shaking at 120 rpm for


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18 h. The number of colony-forming units (CFUs) was determined and compared with control after 12 h incubation at 37 °C.

2.6.4. Fluorescent imaging of bacteria-hydrogel contact on agar gel Luria broth (LB) agar gels were prepared in six-well plates according to the manufacturer’s instructions. A circular piece of the agar gel was removed by 1 ml blue pipette tip to make a cylindrical cavity inside the gel. A freshly prepared hydrogel was transferred into the cavity and incubated at 37 °C for 30 min, after which the surface of sample and agar gel were washed with 1×PBS twice. A volume of diluted bacteria solution (400 µl) in 1×PBS (E. coli, 104 CFU ml-1) was added to each well and rocked to uniformly cover the surfaces. An agar gel with cavity and without hydrogel was used as control. The plate was incubated at 37 °C for 18 h, followed by staining bacteria with green fluorescent SYTO 9 dye (Life Technologies, Singapore). After 15 min, the plate was imaged using Confocal Laser Scanning Microscopy.

2.7. In vitro cytotoxicity of hydrogels The cytotoxicity of the hydrogels was evaluated using hepatocellular carcinoma (HepG2) liver cancer and MDA-MB-231 triple negative breast cancer cell lines, Hela, MRC5, NIH3T3 fibroblast, and smooth muscle cell. The cells were harvested from 25 cm2 T-flask culture dish and seeded at a density of ~4×105 cell per well onto a 24-well plate and allowed to attach overnight. The cells were incubated with freshly prepared hydrogels in 400 µl of DMEM supplemented with 10% FBS and 1% antibiotics at 37 °C and 5% CO2. After 24 h, the cells were rinsed with 1×PBS twice and resuspended in 400 µl fresh growth medium supplemented with 80 µl of CellTiter 96® AQueous One Solution Cell Proliferation Assay (Promega Corporation,



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U.S.A.). After 4h, the absorbance of the solution was measured at 492 nm using TeCAN microplate reader. Cells without hydrogels were used as control. To examine the long-term cytotoxic effect of hydrogels, cells (MDA-MB-231) were incubated with hydrogels (bPEI(1%)/Alg-Ald (3%)) for 1, 2, and 3 days. Then, cells were washed with 1×PBS and stained with fluorescence-based LIVE/DEAD® Cell Viability Assays according to manufacturer’s instructions. The samples were imaged using Confocal Laser Scanning Microscopy.

2.8. Three-dimensional spheroids for cell viability assay The MDA-MB-231 tumor spheroids were prepared according to method reported by Friedrich et al. 35 For homospheroid formation, cell suspension was diluted to a density of 5000 cells ml-1 in high-glucose Dulbecco’s modified Eagle’s medium (DMEM) supplemented with 10% fetal bovine serum (FBS). 200 µl of diluted cells was added to wells of 96-well plate precoated with 50 µl agarose gel (1 wt. %). The spheroids formed after 5 days and subsequently were incubated with freshly prepared microparticles/hydrogel delivery systems (day 0). The size and morphology of the 3D tumor spheroids were monitored at predetermined intervals and compared with control groups.

2.9. Statistical analysis All experiments have been conducted in triplicate or more per condition. The statistical analysis (student t-test) was performed, where p< 0.05 and p < 0.01 were considered statistically significant.


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3. RESULTS AND DISCUSSION 3.1. Particle fabrication and characterization Coaxial electrohydrodynamic atomization (CEHDA) technique was employed in this study for the fabrication of core-shell microparticulate structures. As seen in Figure 1, the size of microparticles increased with the increase of shell flow rate, while the morphology of the particles slightly changed from irregular-shaped microparticles to well-formed microspheres. Our previous studies demonstrated that dichloromethane (DCM) was a suitable candidate for dissolution of PLGA and PDLLA and fabrication of microparticles using CEHDA technique. In this study, a mixture of DCM and dimethyl sulfoxide (DMSO) (5:1 v/v) was used for the core solution due to the higher solubility of Cis-DDP in DMSO compared to aqueous solution. The presence of DMSO in the core solution led to the formation of microparticles with smoother surfaces. However, it adversely affected the morphology of the particles, particularly at a lower shell flow rate (1.0 ml h-1). DCM has a higher vapor pressure compared with common organic solvents such as DMSO and dimethylformamide (DMF). Therefore, micro-droplets can quickly form and solidify after breaking-up of the polymer jet at the tip of the nozzle and flying toward the collecting substrate. Although the fast evaporation of organic solvents is crucial to fabricate microparticles with uniform size distribution, it can affect the surface morphology of particles produced. Solvent evaporation rate affects polymer diffusion, which plays a determinant role in morphology of final products. A greater particle size and higher surface roughness are the major impacts of using a highly volatile solvent. The fast solvent evaporation usually confines the interaction and rearrangement of polymer chains inside a droplet flying toward the oppositely charged substrate. In addition, the fast evaporation of the solvent may cause a highly porous surface, which can



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influence the release rate at a later stage. The presence of DMSO reduces the volatility of DCM and allows polymer chains to re-arrange and form a less porous surface during the solidification step. However, the microparticles may reach the substrate before being fully dried, where the particle’s shape may further change at this stage. As seen in Figure 1, the size of the particles increased as the shell flow rate changed from 1.0 ml h-1 to 2.5 ml h-1. However, the formation of secondary droplets at high flow rate (i.e. 2.5 ml h-1) adversely affected the size distribution of the particles. The size of microparticles is correlated with the flow rates of the core and shell solutions. Therefore, the flow rates must be carefully chosen to allow complete solvent evaporation from droplets before reaching the collector. The core-shell structure of microparticles was confirmed using laser scanning confocal images, where the fluorescent dye (coumarin 6) was mixed with the core polymer solution prior to fabrication process. From the confocal image, the core compartment was completely surrounded by the shell matrix, where the fluorescence intensity profiles showed that the shell layer was uniformly distributed around the core (Figure 2). However, due to the diffusion and dissolution of coumarin 6 into the shell solution containing DCM, the estimation of shell thickness (~ 8-10 µm) using the confocal image may not be so accurate.


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Figure 1. SEM micrographs of electrosprayed microparticles depicting the surface morphology and the size distribution of core-sell microparticles: core flow rate: 0.5 ml h-1and shell flow rate: (a, a’, a’’) 1.0 ml h-1, (b, b’, b’’) 1.5 ml h-1, (c, c’, c’’) 2.0 ml h-1, (d, d’, d’’) 2.5 ml h-1.



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Figure 2. Confocal micrographs illustrating coumarin 6 (green) loaded core-shell microparticles. The radially averaged fluorescence intensity profiles depicting the spatial location of the core compartment inside microparticles (scanning intervals:32).

3.2. CFD simulation of CEHDA process To observe the Taylor cone-jet mode and droplet formation, which dominate the formation of core-shell microparticles, a CFD simulation was carried out near the tip of nozzle. Figure 3


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illustrates the CFD simulation results involving Taylor cone-jet and droplet formation processes, electric potential distribution, electric field strength distributions, charge density, and velocity vector. Two fluid meniscus were set at the tip outlets of the nozzle as initial conditions for both streams. As time proceeded, the fluids were elongated along the axis and formed a cone-jet due to the force balance among the gravity, surface tension, electric stress, and viscous stress. Finally, the jet broke into droplets and formed core-shell droplets with the size of ~30 µm. The final particle size (~13µm) can be predicted from droplet size by Equation S8 which is correlated well with the experimental result. From the velocity vector distributions, the motions of fluids inside the Taylor-cone can be obtained. The liquids moved down along two vortex edges located near liquid-gas interface. At the liquid cone apex, the fluids moved into the jet flow, followed by jet breaking up and droplet formation.

The results also showed that the highest potential

appeared at the tip and liquid cone, while the highest electric field strength was located at the gas-liquid interface near the cone apex. This phenomenon can be attributed to the presence of induced charges at the surface of liquid cone, especially near the cone apex as confirmed by the results of charge density distribution.



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Figure 3. The CFD simulation results: (a) Taylor cone-jet and droplet formation processes (Core: PLA solution, Shell: PLGA solution, surrounding fluid: air), (b) electric potential distribution, electric field strength distribution, charge density distribution, and (c) velocity vector. Operating conditions: Voltage (5600V), shell:core flow rate (2.5:0.5 ml h-1).

3.3. Alginate-aldehyde and hydrogel formation The synthetic procedure of injectable hydrogel was illustrated in Figure 4. Three kinds of alginate-aldehyde of different molecular weight were synthesized using sodium metaperiodate for various duration. Periodate molecules target the α-glycol groups of polysaccharides and form their dialdehyde derivatives. As the aqueous solution of alginate forms a highly viscous solution even at low concentrations, the oxidation of alginate was preferably preformed in a heterogeneous medium composed of ethanol:water (1:1 v/v). As reported by Balakrishnan and Jayakrishnan

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, the oxidation degree of alginate is associated with the quantity of periodate in

the reaction medium and the duration of the reaction. It means that a longer reaction produces more aldehyde groups ready for reacting with cross-linkers at a later stage. However, an extensive depolymerization happens during the oxidation reaction (Table 1), which may adversely affect the formation and stability of hydrogels. Depolymerization reaction can be attributed to the formation of 1-hydroxyethyl radicals produced as a side product during the desired reaction. 36


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Figure 4. Schematic representation of the preparation of injectable hydrogel composed of alginate aldehyde (Alg-Ald) and PEI-25k.



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Table 1. Gel permeation chromatography (GPC) results of sodium alginate and Alg-Ald. Item Sodium Alginate (MMw) Sodium Alginate (HMw) MMw Alg-Ald -6h HMw Alg-Ald -12 h HMw Alg-Ald -18 h

Mn 207,200 259,500 134,000 141,100 126,100

Mw (g mol-1) 223,000 270,300 162,600 166,300 156,600

PDI 1.08 1.04 1.21 1.18 1.24

The hydrogel was formed upon mixing of Alg-Ald solution with PEI-25k solution in deionized water pH~7.5. PEI-25k chains act as a cross-linker where their ε-amino groups reacted with aldehyde groups of oxidized alginate via Schiff’s base formation at physiological temperature. The concentration of Alg-Ald solution was kept constant at 3 wt. % throughout the experiments, while the PEI concentration was changed from 1 wt. % to 10 wt. %. The time required for hydrogel formation showed a PEI concentration dependent trend. The time was quite short when the PEI concentration was 1 wt. %, while no obvious gel network was formed at PEI concentrations above 5 wt. %, even after 15 min. This was due to the change of primary amines concentrations available to react with aldehyde (RCHO) groups. At low PEI concentrations (>5 wt. %), primary amines of one PEI molecule react with a number of RCHO groups from different alginate chains. Therefore, it can make a dense cross-linked network. However, as PEI concentration increases, the probability of the reactions between amines of one PEI molecule and RCHO decreases. It can be attributed to the physical mixing ability of the high concentration of bPEI with Alg-Ald solution. These solutions cannot be mixed homogeneously, where the access of RCHO groups to primary amines suddenly reduces. Thus, a weak polymer network with a non-porous morphology is formed (Figure S1).

3.4. Chemical characterization of hydrogel


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FTIR spectra of sodium alginate, oxidized alginate, and composite hydrogel are shown in Figure 5a. The spectrum of sodium alginate showed characteristic absorption bands around 945 cm˗1 (C‒O stretching), 1029 cm˗1 (C‒O‒C stretching), 1130 cm˗1 (C‒C stretching), and 1321 cm˗1 (C‒O stretching). The strong bands at 1612 and 1413 cm˗1 are associated with asymmetric and symmetric stretching bands of the carboxylate salt groups. The appearance of aldehyde symmetric band at 1724 cm˗1 confirmed the formation of alginate-aldehyde due to reaction with sodium meta-periodate. However, hemiacetal formation of free aldehydes, in some cases, may reduce the intensity of this band. As seen in Figure 5b, the aldehyde symmetric band at 1724 cm˗1 disappeared upon the mixing of alginate-aldehyde and PEI and formation of hydrogel. Therefore, the absence of the band in hydrogel spectrum clearly indicated the formation of Schiff’s base (i.e. C=N chemical bond) between ‒CHO and ‒NH2 groups. Figure 5c presents the solid-state

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C-NMR spectrum of the hydrogel upon mixing alginate aldehyde and PEI-25k

solutions at room temperature. Compared to bare sodium alginate spectrum shown in Figure S2, the peaks appeared at ~35-60 ppm and 160-163 ppm are assigned to PEI carbons and Schiff base carbon (-HC=NH-), respectively. The presence of -HC=NH- peak in the hydrogel

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C-NMR

spectrum is in a good agreement with the results of FTIR analysis, where the aldehyde peak appeared and disappeared upon the oxidation of alginate molecules with NaIO4 and formation of the hydrogel, respectively. The elemental analysis (EA) results are presented in Table S2, where the molar ratio of carbon/nitrogen ⁄ is calculated from the weight percentage of both components. Due to constant carbon: nitrogen ratio in each PEI unit (2C:1N) and alginate unit (6C:0N), the grafting rate can subsequently be calculated using Eq. S9. Using this equation, α= 0.003 for PEI (1%)/Alg-Ald (3%) is the maximum grafting ratio.



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The data provided by TGA analysis of sodium alginate, alginate aldehyde, and the hydrogel were presented in Figure 5d and Figure S3. The thermograms showed different thermal behavior of the samples. Sodium alginate exhibited three degradation steps, where the first one occurred ~100 °C. The first step of mass loss corresponded to the elimination of free water and bound water from the sample. The second step occurred ~220 °C that can be attributed to the decomposition of alginate at C‒H and C‒O‒C bonds in the main polysaccharide chain and vaporization and elimination of volatile products. The third step was associated with the formation of Na2CO3 and other carbonized materials. The thermogram of alginate-aldehyde, however, showed a few differences in the peak area and position as compared with bare alginate. The first degradation step of alginate-aldehyde was faster than that of alginate. The variation on the peak area and position during the first step of mass loss was corresponded to water molecule interactions with polymer chains and the lower capacity of alginate-aldehyde for holding of water. In addition, the second degradation step of alginate-aldehyde exhibited a maximum peak at ~200°C, which was slightly lower than the maximum peak (~220 °C) of alginate thermogram. As the stability of the samples is directly correlated with their decomposition behavior, the results demonstrated a significant decrease in the thermal stability of alginate-aldehyde compared to pure alginate. In the case of hydrogel, the results demonstrated a two-step mass loss similar to that of alginate-aldehyde, followed by a gradual mass loss from ~220 °C onward. The new peak appeared at ~440 °C might be due to the decomposition of the materials and was in agreement with results (i.e. a less thermal stability for biopolymeric Schiff based materials) of previous study.37 The data provided by DSC curves (Figure 5e) were in agreement with TGA results. The DSC curves showed that dehydration process of pure alginate occurred at a lower temperature as


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compared with other samples. The exothermic peaks ~200 °C and 220 °C can be attributed to decomposition of the samples, where a broad peak in this area was associated with a gradual decomposition of hydrogel due to temperature increase.

Figure 5. Chemical characterization of synthetic hydrogels. (a) and (b) FTIR spectra of the hydrogel formed after the mixing of alginate-aldehyde with PEI at different concentrations. (c) Solid state

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C-NMR spectrum of hydrogel. (d) Thermogravimetric analysis of bare sodium

alginate, oxidized alginate, and synthetic hydrogel. (e) DSC curves of bare sodium alginate, oxidized alginate, and synthetic hydrogel.



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3.5. Preparation of microparticles embedded hydrogel Figure S4 shows the morphological images of core-shell microparticles embedded in AlgAld/PEI hydrogel networks. Microparticles were dispersed in Alg-Ald (3 wt. %) solution and PEI-25k (1 wt. %) solution was subsequently added to form the hydrogel matrix. Figure S1a shows the morphology of the hydrogel prior to incorporation of microparticles. The hydrogel had a highly porous structure, which can significantly help to absorb large amount of water and increase its biocompatibility. The addition of microparticles (25, 50, 75 w/w %) did not change the morphology of hydrogel and microparticles (Figure S4 b-d). However, due to hydrophobicity of microparticles, the homogeneity of final product is affected by the increase of particle weight percentage. This conclusion was further proved using higher magnification SEM micrographs for the particles embedded hydrogels at 50 w/w % and 75 w/w % (Figure S4 e-f). The dispersion of microparticles (Figure S4-g) within the hydrogel was further examined using fluorescence dyeloaded microparticles and corresponding fluorescence intensity profiles (Figure S4-h). Although the intensity profiles correlated well with the concentration of microparticles, a low light intensity was observed near the gel surface in tube (iv), which means that the particles could not be uniformly dispersed at high (~ 75 w/w %) concentration.

3.6. Physical characterization of hydrogel The swelling ratio (SR) of hydrogels with different compositions was measured in 1×PBS buffer at different pH values and 37°C (Figure 6). The hydrogels generally reached to a state of swelling equilibrium within half an hour, although average quantities were slightly fluctuating around the equilibrium ratios. As illustrated in Figure 6a, the SR of the hydrogels increased significantly with the decrease of PEI concentration used for hydrogel preparation. However, the


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SR of hydrogels cross-linked by PEI at concentrations of 1.0 and 5.0 wt. % fluctuated around 1800 % of their initial dry weights and remained almost constant for 12 h. The water absorption in the hydrogels is a function many of variables such as crosslink density, hydrogel microstructure (porous or poreless), network parameters, nature of the materials etc. For instance, the crosslink density determines the distance between two crosslink nodes on a single polymer chain. Therefore, a longer distance results in a lower crosslink density. In addition, the swelling process can be controlled by diffusion and relaxation processes. The diffusion process, identified as the rate-determining step at the beginning of the swelling process, depends on solvent molecular weight, porosity of hydrogel, and temperature. The chain relaxation step occurs slower than diffusion process at low crosslink density. Therefore, as explained by Omidian and Park 38, the chain relaxation step is considerably slower than diffusion process at a high crosslink density, which results in a sharp increase in SR, followed by an equilibrium step. The effect of alginate molecular weight on SR was investigated as shown in Figure 6b. The concentration of PEI solutions was kept constant at 1 wt. % for all experiments. As can be seen, the SR quickly reached an equilibrium condition for hydrogel prepared using medium molecular weight alginate (MMw 6), while it was gradually increasing for high molecular weight alginate oxidized for 12 h and 18 h. However, the results showed that SR could slightly decrease when the oxidation process was longer (i.e 18 h). The longer the oxidation time, the more aldehyde groups for crosslinking process. As a result, the higher crosslink density confined the polymer relaxation steps and reduced SR. On the other hand, longer oxidation period caused polymer decomposition and reduced molecular weight. The lower the molecular weight, the more the compacted gel network is and consequently the lower SR.



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The pH of the aqueous medium had a marginal effect on SR and became effect-less after ~ 1 h (Figure 6c). pH can partially increase the protonation of amine groups available in the solution, which may delay the crosslinking process and provide a slightly longer relaxation period for polymer chains inside the gel network. These results were ratified by PEI release test at different pH within 5 days (Figure 6d), where the amount of PEI release in the first day was negligible for all pH. However, the carbonyl-amine bonds showed greater sensitivity to pH less than 6 in the following days.

Figure 6. Swelling ratios of different hydrogels as a function of time: (a) hydrogels prepared at different Ald-Alg: PEI weight ratios, (b) Hydrogel prepared at Ald-Alg: PEI of 3:1 w/w % where high molecular weight alginate was oxidized by NaIO4 for 12 h and 18 h. Medium molecular weight alginate was oxidized for 6 h using the same process. (c) The swelling ratios of hydrogels


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at different pH values. (d) PEI release percentage during 5 days at different pH values (hydrogel: Ald-Alg: bPEI of 3:1 w/w %). All data are expressed as the average± standard deviation of three independent experiments. (The deswelling behavior of the hydrogel is reported in supporting information).

The degradation of hydrogel samples was investigated in vitro and the results were presented in Figure 7. Among different hydrogels, bPEI (1%)+Alg-Ald (3%) was the most stable sample over 60 days, while others lost their weight considerably faster over the same period. As seen in Figure 7, the initial concentration of bPEI solution strongly affects the stability of the hydrogels and degradation rates, where bPEI (10%)+Alg-Ald (3%) that did not form obvious gel completely disappeared before day 30.

Figure 7. Hydrogel degradation results. All data are expressed as the average± standard deviation of three independent experiments.

The dynamic mechanical behavior of bPEI (1%)/Alg-Ald (3%) hydrogel was investigated by oscillatory rheology at 37 °C. First, the stress sweep at a constant frequency was conducted to identify the linear viscoelastic region (LVR) profiles of the hydrogel. The effect of the stress 31 ACS Paragon Plus Environment


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amplitude on the cross-linked hydrogels is shown in Figure 8a. Over the linear range, the applied stress results in strain. However, beyond LVR, Gʹ (elastic modulus) sharply decreases due to the structure breakdown upon increasing the deformation imposed. Therefore, the pressure (~200 Pa) at which the hydrogel network begins its nonlinear viscoelastic behavior was determined as critical shear stress. As the linear part happened over the range of 0.5 to 200 Pa (Figure 8a), the frequency sweep experiment was subsequently conducted at 10 Pa to investigate the stability of three-dimensional cross-linked gel networks over a frequency range from ~0.6 to 600 rad s˗1. As shown in Figure 8b, there is a large viscoelastic plateau between ~0.6 and 200 rad s˗1, which is an indication of a stable hydrogel structure. No cross-point between Gʹ and Gʺ was also observed, hence representing the stability of the hydrogel structure. In addition, the frequency independent behavior of elastic modulus indicates the solid-like behavior of the hydrogel. At low frequency, the polymer chains between the cross-link points can rearrange themselves due to comparable relaxation time scales, while at higher frequencies, they are impotent to rearrange themselves within a short period of imposed motion (chain stiffening) and show solid-like behavior.


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Figure 8. Rheological analysis of bPEI (1%)/Alg-Ald (3%) hydrogel. (a) Stress sweep and (b) frequency sweep.

3.7. Antibacterial activity of hydrogels To evaluate the antibacterial property of the hydrogels, the surface of the hydrogels was challenged with model Gram-negative (E.coli) and Gram-positive (B. pumilus) bacteria. Figure 9a and 10c demonstrate the results of antibacterial tests in which the surface of hydrogels prepared with varying amount of PEI was incubated to 104 CFU ml-1 of bacteria. The empty tissue culture wells and the same amount of PEI used for the preparation of hydrogels were considered as negative and positive controls, respectively. The antibacterial activity of formulations as well as controls were evaluated by the colony formation ability within 18 h postexperiment times. The results showed that antibacterial activity of hydrogels increases with the increase of PEI concentration in the formulations. On the other hand, a bare 3 wt. % alginate solution show no antibacterial activity against both E.coli and B. pumilus, similar to empty wells (negative controls). For instance, at 1 wt. % PEI the hydrogel showed ~100 times greater antibacterial activity compared with bare alginate. The inhibitory effect of hydrogels was significantly higher at PEI concentration ≥ 2.5 wt. %, where no colony was formed after 18 h incubation on agar plates. These results suggest that the protonated amines in PEI molecules can directly influence its antibacterial behavior. Figure 9b and 9d presented the colony formation ability of bacteria as a function of increasing initial CFU for a hydrogel prepared from 3 wt. % alginate+1 wt. % PEI. The data showed that the hydrogel is highly active at low bacteria concentrations. In addition, a significant reduction in bacteria activity (~1000 times) was



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observed at extremely high bacteria concentration (108 CFU ml-1). This high concentration was ~10 million times greater than that is expected in a contaminated operating theatre. 39-40


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Figure 9. Antibacterial activity of hydrogels: Alg-Ald cross-linked with different concentration of PEI-25K were incubated with 104 CFU ml-1 (a) E.coli or (c) B. pumilus for 18 h. Hydrogel surface (Alg-Ald: PEI 3:1 w/w) was challenged with serially diluted (b) E.coli or (d) B. pumilus in PBS for 18 h. (e) The time course of bacterial (E.coli) killing experiment. (f) Contactdependent antibacterial effect. The control experiment (purple bars) was performed by incubating bacteria (104 CFU ml-1) without hydrogels for the corresponding time points.

Figure 9e shows the time-dependent antibacterial test of the hydrogel prepared from 1 wt. % PEI. The colony formation ability of bacteria incubated with the hydrogel at various time-points was evaluated and the data showed a time-dependent activity where the number of colonies gradually decreased as the incubation period extended. All data presented above suggests a contact-dependent mechanism of action for antibacterial activity of the hydrogels. To assess this hypothesis a set of quantitative and qualitative experiments was performed in vitro. The hydrogel prepared was loaded into a transwell tissue culture insert and the insert was placed in the culture plate containing bacteria in 1×PBS, where the porous membrane of the insert separated the hydrogel and bacteria. After 18 h, the effect of PEI molecules leached from the hydrogel on the activity of bacteria was determined. The results of bacteria-hydrogel contact test were presented in Figure 9f. As can be seen, the number of colonies formed after the experiment was very close to that of control groups. The data suggest that most of PEI molecules reacted with Alg-ald molecules and did not leach from the hydrogel as previously indicated in Figure 6d (PEI concentration in supernatant). This mechanism was qualitatively investigated by bacteria-hydrogel contact experiment on agar plate. A cylindrical portion of preformed agar gel was removed and the hole was filled with the hydrogels. Next, the


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surfaces of hydrogel and agar gel were exposed to bacteria solution at 104 CFU ml-1 for 18 h. Figure 10 shows both the bright field and fluorescent image of bacteria stained with SYTO-9 fluorescent dye. The lack of green fluorescence observed on the area occupied by the hydrogel indicated that there were no viable bacteria on the hydrogel surface. However, bacteria could proliferate on the surrounding agar surface. Based on these results, the hydrogel is a contactedbased bactericidal surface, which kills adhering bacteria.

Figure 10. The viability of bacteria (E.coli) on the surface of hydrogel. E.coli viability was imaged using confocal microscopy after 18 h of incubation on agar gel/hydrogel interface. (a) The bright field, (b) the fluorescence micrograph of bacteria stained with SYTO 9, and (c) the overlay image of bacteria and gels. The white dash-line indicates the interface of agar gel/hydrogel. Scale bars: 100 µm.

3.8. Biocompatibility of hydrogels Since biocompatibility is a crucial factor for their successful applications of synthetic biomaterials, the cytocompatibility of the hydrogels was evaluated by MTS assay using six different cell lines. The cell lines were selected from different tissues, where the drug delivery



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system proposed in this article may be used. Among four different hydrogels tested for cytocompatibility, the hydrogel with the lowest PEI concentration (i.e. Alg-ald:PEI 3:1 wt. %) showed minimum toxicity and maximum cell viability (Figure 11a). However, the number of live cells was gradually decreased with increasing PEI concentration from 2.5 to 10 wt. %. Notably, cell proliferation assay revealed that the hydrogel prepared with 1 wt. % of PEI should be more stable than other hydrogels due to formation of a well- interconnected network which prevents the premature release of non-crosslinked toxic PEI. A similar trend was found for the cell lines incubated with medium molecular weight and high molecular weight hydrogel at various PEI concentrations (Figure 11 b-c). Therefore, it can be concluded that molecular weight of the oxidized alginate has no detrimental effect on cell growth and proliferation. The extended cell viability test was performed for three days. Cells were stained using LIVE/DEAD assay and observed using laser scanning confocal microscopy (Figure 11d). The results showed a few dead cells (red color), and cells maintained normal cell morphology for three days. These results demonstrated that the hydrogel (i.e. Alg-ald:PEI 3:1 w/w %) is an appropriate candidate for longterm drug delivery applications.


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Figure 11. Results of in vitro biocompatibility tests for cells co-cultured with hydrogels. (a) Cell viability percentage after 24 h incubation with hydrogels at various PEI concentrations. Note: Smooth muscle cells were cultivated in smooth muscle cell medium (SMCM) supplemented with 2% of FBS, 1% of cell growth supplement, and 1% of penicillin-streptomycin antibiotic solution. The viability percentage of (b) HepG2 and (c) MDA-MB-231 cells after 24 h treatment with 39 ACS Paragon Plus Environment


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hydrogels at various PEI concentrations. (d) Live/dead cell viability staining images of MDAMB-231 cells after three consecutive days co-cultured with hydrogel (Alg-ald:PEI 3:1 w/w %). The green color demonstrates live cells and the red color demonstrates dead cells. Dead cells were indicated by red dash-circles. All data represent as mean ± S.D. of n=4 independent experiments. Scale bars: 200 µm. 3.9. In vitro drug release: microparticles and microparticles-embedded hydrogels The molecular weight of an anti-cancer agent and its affinity for the carrier’s molecules are important parameters contributing to the design of a successful controlled release system. Thus, the ability of microparticles for the sustained release of verapamil (VRP), paclitaxel (PTX), and cisplatin (Cis-DDP) as model drugs was examined in vitro. Table S3 presents the results of drug loading content (DLC) and encapsulation efficiency (EE) of the particles. The release profiles of both formulations showed a small initial burst of ~11% VRP and ~12% PTX for formulation A1, and ~13% VRP and ~4% PTX for formulation B1 (Figure S6). However, after an initial premature release, the drugs were gradually released in a sustained manner over a period of 60 days, indicating the progressive degradation of polymeric microspheres for a sustained and longterm release. In formulation A1, both drugs exhibited a near zero order release rate, however, for VRP in formulation B1 (Figure S6-b), it showed a sharp increase in release profile after 30 days. In addition, for formulation A1 (Figure S6-a), both PTX and VRP release curves overlapped with each other in the first 10 days, after which VRP was released slightly faster, while the gap between the two release curves was significantly larger in formulation B1 (Figure S6-b). Moreover, the slope of VRP curve suddenly increased after 30 days, causing an even greater deviation from PTX curve. Therefore, VRP was released significantly faster and in a larger dosage, making it seemed to be released first while PTX was released second. These results


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indicated that highly versatile drug carriers could be designed by simply swapping the drugs in a one-step CEHDA fabrication technique to present a parallel or a sequential release of drugs. This flexibility is an advantage of the CEHDA fabrication technique compared to other techniques such as the emulsion/ solvent evaporation method.

41-42

Emulsion/ solvent evaporation method

involves a rapid stirring step that adversely promotes the disintegration of the drugs and confines loading capacity of hydrophilic drugs due to prolonged contact with aqueous environment. In the case of Cis-DDP and PTX loaded microparticles (Table S4), the release behavior was considerably different from what was previously mentioned above. For the case of Cis-DDP dissolved in water, the release of Cis-DDP from both formulation B2 and C2 (Figure 12) displayed an initial burst of approximately 40%, followed by a sustained release throughout a period of 45 days in a near zero-order rate. This initial burst release can be attributed to the premature release of drug molecules from the surface of particles and/or the diffusion of drugs placed close to the surface. Furthermore, the hydrophilic nature of Cis-DDP may contribute to higher premature release, as it is more readily available for release into surrounding medium. In contrast, the release of PTX from both formulation A2 and C2 did not show significant initial burst, but just a progressive and sustained release throughout the study. The steady release of PTX could be explained by the drug diffusion within the microspheres, hence giving a predictable trend. As time proceeded, the continuous penetration of aqueous medium caused polymer degradation, pore formation, thinning of the shell layer

43

, and consequently faster

release of drug into surrounding medium. As the limited dissolution of Cis-DDP in water (25 mg ml-1). In this case, the drug loading content was significantly improved (Table



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S4). However, a huge initial burst release of Cis-DDP (~82% for Formulation B3 and ~78% for Formulation C3) was observed within the first day of experiments (Figure 12d). It was due to the miscibility of DMSO and DCM that led to the undesirable diffusion of core solution into shell counterpart during particle fabrication, where the effective thickness of the shell compartment markedly decreased. Therefore, Cis-DDP could readily pass the shell barrier and dissolve into the surrounding medium. In order to control the initial burst of Cis-DDP, the previously introduced hydrogel was employed to entrap the drug-loaded microspheres. Thus, the initial burst was significantly reduced to ~40% (Figure 12d) due to a more complex diffusion pathway to be passed. The hydrogel matrix functioned as an additional barrier to restrict direct exposure of microparticles to PBS medium, thereby, reducing its high initial burst as well as making its release more sustainable over a longer period. As shown in Figure 12e, after preparation of composite formulation, 75 w/w % microspheres-loaded hydrogel exhibited a slightly lower initial burst and cumulative release as compared to 25 w/w% and 50 w/w% composite formulations, though all three release curves had very similar shape and gradient. The composite formulation C3 showed similar trend for CisDDP, while release rate of PTX was undoubtedly reduced (Figure 12f). Regardless of the initial burst release of Cis-DDP, the new system showed a simultaneous release of both agents in a controlled and sustained manner, here is highly promising for the synergistic treatment of tumors such as triple-negative breast cancer (TNBC).


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Figure 12. The release profiles of Cis-DDP and PTX from core-shell microparticles: (a) the scheme of microparticles loaded with Cis-DDP and PTX (formulation A2: PTX (shell), formulation B2: Cis-DDP (core), formulation C2: PTX (shell) and Cis-DDP (core)); (b) the release profiles of formulation A2 and B2 (Cis-DDP was dissolved in water); (c) the release profile of formulation C2 (Cis-DDP was dissolved in water); (d) the initial step of Cis-DDP



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release profile from formulation B3 and C3 (Cis-DDP was dissolved in DMSO) with and without hydrogel; (e) the release profile of formulation B3 (Cis-DDP was dissolved in DMSO) and hydrogel at 25, 50, and 75 w/w % loading ratios; (f) the release profile of formulation C3 (CisDDP was dissolved in DMSO) and hydrogel at 50 w/w % loading ratio. pH of PEI-25k and AlgAld solutions were adjusted at physiological pH (pH~7.0) using hydrochloric acid (32.5%) prior to hydrogel formation. Data represents average ± standard deviation of three independent measurements, n = 3.

3.10. Drug release effect on 3D MDA-MB-231 spheroids Although the traditional 2D cell-culture is suitable for mechanism study and rigorously control the experimental and treatment conditions, it usually provides overestimated or underestimated results that may not be reproduced in vivo. Therefore, a 3D tumor spheroid model is employed in the current study as it demonstrates dose-response relationships that may better predict tumor responses in vivo in future studies. MDA-MB-231 spheroids in a prolonged study of 10 days were used to evaluate the efficacy of combined formulation released from microparticles/hydrogel composite. As shown in Figure 13, the composite formulation exerted a greater inhibitory effect on the growth of tumor spheroids, as compared with control and free drug groups. The free drug treatment temporarily inhibited the growth of spheroids and the size of the spheroids remained constant or increased within 10 days. In contrast, the spheroids slowly responded to the treatment of composite formulation at the early stage, while the responses were pronounced over 10 days and significantly inhibited the growth of spheroids (Figure 13b and 13c). The fast but short period response of spheroids to free drug treatment was due to exposure with high concentration of drugs. However, to mimic the in vivo conditions, concentration of


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drugs was gradually reduced over 48 h. Therefore, the spheroid could restore their proliferation conditions after ~8 days. Despite the free drug results, higher drug efficacy observed for spheroids incubated with the composite formulation could be attributed to the controlled and sustained release of Cis-DDP and PTX from the new delivery systems. Although the initial drug concentrations were not sufficient to cause significant cell death, the gradual accumulation of drug within the spheroid killed cells in the outer rim of the spheroid, where it promoted penetration of drugs into the primed tumor and caused a substantial cell death.

Figure 13. (a) Bright-field images of tumor spheroids treated with combination of Cis-DDP and PTX: free drug (second column) and composite formulation (third column) (Scale bar: 200µm). (b) Growth of tumor cell spheroids upon exposure to microparticle formulations over 10 days. (c) cell viability of tumor spheroids upon exposure to microparticle formulations at day 10. Data represents average ± standard deviation of three (n = 3) independent measurements.

4. CONCLUSIONS



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Cis-DDP and PTX loaded core-shell microparticles were fabricated using the CEHDA techniques. A CFD simulation model was developed and utilized for the optimization of CEHDA operating parameters prior to fabrication of microparticles. The Alg-ald:PEI hydrogel was successfully synthesized and evaluated for entrapment of microparticles inside a porous matrix, modulating release profiles, and antibacterial activity against Gram-positive and Gram-negative bacterial via a contact-dependent mechanism. Moreover, three-dimensional MDA-MB-231 spheroid studies demonstrated that the combination of Cis-DDP and PTX released from the proposed microparticles/hydrogel formulations could provide greater efficacy and gradually reduce the size of spheroids. In contrast, the free drug treatment was impotent to effectively reduce cell viability in tumor spheroid in vitro. The new composite formulation may hold great promise for practical applications in cancer chemotherapy through localized delivery of multiple agents in a controlled and sustained manner.

ASSOCIATED CONTENT Supporting Information The Supporting Information is available free of charge on the ACS Publications website at DOI: Tables: CFD model description and governing equations; Drug loading content and encapsulation efficiency of VRP and PTX loaded microparticles; Drug loading content and encapsulation efficiency of Cis-DDP and PTX loaded microparticles. Figures: Morphology of injectable hydrogel composed of Alg-Ald and PEI-25k at various concentrations; Solid-state 13CNMR spectrum of bare sodium alginate; TGA and DTG curves of the hydrogel; The deswelling kinetics of the hydrogel; SEM micrographs of hydrogel morphologies before and after loading with core-shell microparticles; TGA and DTG curves of the hydrogel at different heating rates; Weight changes of the hydrogel over time; In vitro release profiles of PTX and VRP from double-walled microparticles.


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AUTHOR INFORMATION Corresponding Author E-mail: [email protected]; Tel.: +65-65165079; Fax: +65-67791936 Notes: The authors declare no competing financial interest.

Acknowledgements This work is financially supported by A*STAR and National University of Singapore under the project/grant numbers APG2013/40A (A*STAR BMRC Strategic Positioning Fund, A*STAR-P&G Collaboration, R279-000-487-305) and R261-509-001-646 (3D Printing Initiatives), respectively. Pooya Davoodi greatly appreciates the National University of Singapore for PhD graduate research scholarship. We thank Dr. Marc V Garland and Mrs. Li Li Ong for their help with solid-state 13C-NMR experiments.

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Table of Content


An injectable thermosensitive polymeric hydrogel for sustained release of Avastin® to treat posterior segment disease.

Sustained Release of Antibacterial Agents from Doped Halloysite Nanotubes.

A repertoire of peptide tags for controlled drug release from injectable noncovalent hydrogel.

Hyperbranched phosphoramidate-hyaluronan hybrid: a reduction-sensitive injectable hydrogel for controlled protein release.

An injectable sustained release fertility control system.

Sustained and controlled release of lipophilic drugs from a self-assembling amphiphilic peptide hydrogel.

Controlled release of protein from biodegradable multi-sensitive injectable poly(ether-urethane) hydrogel.

Nanoenhanced hydrogel system with sustained release capabilities.

Injectable Chemically Crosslinked Hydrogel for the Controlled Release of Bevacizumab in Vitreous: A 6-Month In Vivo Study.

Biodegradable injectable in situ implants and microparticles for sustained release of montelukast: in vitro release, pharmacokinetics, and stability.

An injectable, dual pH and oxidation-responsive supramolecular hydrogel for controlled dual drug delivery.

PLGA microspheres for onychomycosis treatment.

Injectable small molecule hydrogel as a potential nanocarrier for localized and sustained in vivo delivery of doxorubicin.

Injectable and Self-Healing Carbohydrate-Based Hydrogel for Cell Encapsulation.

Assembly of an injectable noncytotoxic peptide-based hydrogelator for sustained release of drugs.

Intravaginal flux controlled pump for sustained release of macromolecules.

Controlled drug release from hydrogel-based matrices: Experiments and modeling.

Modular approach for bimodal antibacterial surfaces combining photo-switchable activity and sustained biocidal release.

Sustained release and enhanced bioavailability of injectable scutellarin-loaded bovine serum albumin nanoparticles.

Lactation, health, and reproduction of dairy cows receiving daily injectable or sustained-release somatotropin.

Engineering Biomaterial-Drug Conjugates for Local and Sustained Chemotherapeutic Delivery.

Thermoresponsive magnetic nanoparticle--aminated guar gum hydrogel system for sustained release of doxorubicin hydrochloride.

Interactions of chemotherapeutic agents and radiation.

Blend Hydrogel Microspheres of Carboxymethyl Chitosan and Gelatin for the Controlled Release of 5-Fluorouracil.