Reloadable Silk-Hydrogel Hybrid Scaffolds for Sustained and

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Reloadable Silk-Hydrogel Hybrid Scaffolds for Sustained and Targeted Delivery of Molecules Saket Kumar Singh, Bibhas Kumar Bhunia, Nandana Bhardwaj, Sween Gilotra, and Biman B. Mandal Mol. Pharmaceutics, Just Accepted Manuscript • DOI: 10.1021/acs.molpharmaceut.6b00672 • Publication Date (Web): 26 Oct 2016 Downloaded from http://pubs.acs.org on October 28, 2016

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Molecular Pharmaceutics

Reloadable Silk-Hydrogel Hybrid Scaffolds for Sustained and Targeted Delivery of Molecules Saket Kumar Singh1, Bibhas Kumar Bhunia1, Nandana Bhardwaj2, Sween Gilotra1, Biman B. Mandal1* 1

Biomaterial and Tissue Engineering Laboratory, Department of Biosciences and

Bioengineering, Indian Institute of Technology Guwahati, Guwahati – 781 039, India 2

Biological and Chemical Sciences Section, Life Sciences Division, Institute of Advanced

Study in Science and Technology, Guwahati – 781 035, India

*Corresponding author, E-mail: [email protected]; [email protected] Tel: +91-361-258 2225

1 ACS Paragon Plus Environment


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ABSTRACT: Tunable repeated drug administration is often inevitable in number of pathological cases. Reloadable 3D matrices for sustained drug delivery are predicted as a prospective avenue to realize this objective. This study was directed towards sonication-induced fabrication of novel reloadable Bombyx mori silk fibroin (SF) (4, 6 and 8 wt%) hydrogel, injected within 3D porous (8 wt%) scaffolds. The focus was to develop a dual-barrier reloadable depot system for sustained molecular cargo-release. Both the varying SF concentration (4, 6 and 8 wt%) and the sonication time (30, 45 and 60 s) dictated the extent of cross-linking, β-sheet content and porosity (1-10 µm) influencing the release behavior of model molecules. Release studies of model molecules (trypan blue, TB, 961 Da and bovine serum albumin, BSA, 66 kDa) for 28 days attested that the variations in their molecular weight, the matrix crosslinking density and the scaffold-hydrogel interactions dictated the release behavior. Ritger and Peppas equation was further fitted into the release behavior of model molecules from various SF matrices. The hybrid constructs exhibited high compressive strength along with in vitro compatibility using primary porcine chondrocytes and tunable enzymatic degradation as assessed for 28 days. The aptness of the constructs was evinced as a reloadable model molecule (BSA and Fluorescein isothiocyanate-Inulin, 3.9 kDa) depot system through UVvisible and fluorescence spectroscopic analyses. The novel affordable platform developed using silk scaffold-hydrogel hybrid constructs could serve as sustained and reloadable drug depot system for administration of multiple and repeated drugs. KEYWORDS: silk, hydrogel, scaffold, reloadable, drug delivery, tissue engineering


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1. INTRODUCTION Most of the current conventional drug delivery implants used in clinic is of single use with no ability to reload once the drug is exhausted. Drug eluting scaffold platforms are desired clinically for treatment of various chronic conditions, wherein a spatiotemporal release of the drug is needed coupled with the feasibility to refill this depot in a minimally invasive manner. Use of such reloadable platforms would find tremendous implications in preventing frequent reoccurrence of restenosis, treatment of cancer and its recurrence management, curative stimulation of angiogenesis in order to re-establish microvasculature in ischemic condition.1,2 Similarly, for non-infectious intermediate or posterior uveitis, macular edema and age-related macular degeneration due to diabetic retinopathy, a repeated local administration of drug is a requisite.3 Systemically administered drugs must be given in high doses in order to reach an effective concentration thereby developing side effects. Similarly, repeated drug administration through injection in a single place is subjected to pain and causes local erythema, edema or finally may lead to the infections. Thus, it is essential to develop drug delivery systems which can provide controlled release of drug for potentially longer periods of time. Ideal candidates for these system aim to provide a depot for a repeated and multiple drug administration to the target site. They are advantageous in terms of providing tunable drug release kinetics for an extended period of time and sustained local delivery enabling control over drug availability and minimizing the toxic effects of the drug.2 The biodegradability of these systems ensures relief from complexities due to surgical removal of the implants. The rate of drug absorption depends on the rate of drug release from a formulation.4 This implies that the slow and constant release rate of the drug over a period of time will lead to slow and relatively constant rate of drug absorption over the period of its release. This allows less frequent administration of large doses of drug thereby decreasing the



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fluctuations of the drug concentration in the plasma. These systems are ideal for drugs that have relatively short half-life.4 The use of polymers as carriers is becoming increasingly important in the field of drug delivery. Specifically, biodegradable polymers are widely used because of their ability to be degraded at constant rate matching that of drug release.5,6 The release mechanism of the drugs from these polymers is by swelling, diffusion and degradation. Natural polymers are also used for drug delivery at predetermined rates. Silk is a naturally occurring polymer produced by various organisms such as insect, spider and silkworm species.7 It contains two types of proteins namely, fibroin and sericin which is a water soluble component of silk.8 Silk fibroin can be used for making various types of matrices including hydrogels, scaffolds, microsphere, micro- to nanoparticles, electrospun mats and films in regenerative medicine.912

Silk fibroin (SF) without sericin acts as a good polymer for preparation of matrices because

of its very low inflammatory response with remarkable mechanical properties.13 Mechanical properties of the matrix can play a vital role in controlling various cellular activities besides regulating drug release behavior. Several factors like SF concentration, crosslinking density, type of blending biomaterial and their concentration may influence the biophysical properties of

fabricated

scaffold

and

hydrogel.

Characteristics

such

as

biodegradability,

biocompatibility, tunable porosity, swelling behavior and non-immunogenicity make SF an ideal candidate for drug delivery applications.9,14-16 SF hydrogel and scaffolds have already been used as matrices for drug delivery and tissue engineering application with certain limitations of each.17-19 Hydrogel contains more than 90% of water in bound, unbound and intermediate forms and it regulates various biophysical, enzymatic and drug releasing activities of the matrices.20 These SF hydrogels can be prepared using different methods such as sonication induced gelation,21 vortex induced gelation22 and shear induced gelation processes.23 Some other methods of hydrogel formation include alcohol mediated gelation,24 4 ACS Paragon Plus Environment

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citric acid mediated gelation,25 carbon dioxide mediated gelation,18 acetone induced gelation26 and electrically mediated gelation processes.27 On the other hand, SF scaffold is a three dimensional porous structure which can be prepared by various methods including salt leaching,28 freeze drying,9 freeze gelation,29 water annealing,30 and hexafluoroisopropanol (HFIP) derived silk fibroin scaffold.31 SF matrices have been a popular choice since long in various drug delivery and tissue engineering applications. Some of these applications are matrix mediated gene delivery,32,33 delivery of basic fibroblast growth factor (bFGF),34 transforming growth factor (TGF),35 and platelets derived growth factor (PDGF).36 Some other applications include the encapsulation of human bone mesenchymal stem cell37 or in different types of tissue engineering like connective tissue engineering,38 vascular tissue engineering,39 cartilage tissue engineering,40,41 meniscus tissue engineering42 and intervertebral disc tissue engineering.43 Despite their varied applications, SF hydrogel matrices alone have limitations in providing good mechanical strength along with maintaining optimum biological activity of the impregnated molecules. Along these lines, this study focused on preparing the hydrogel embedded scaffold matrices that are meant to overcome these limitations. In these hydrogelscaffold hybrid constructs, scaffolds provided strength due to their robust nature while hydrogels added soft tissue characteristic holding drug molecules within their fluid environment. The scaffold-hydrogel hybrid system acted as a depot system for drug delivery applications. Model molecules were encapsulated in this system and further their release behavior was studied. This study used two model molecules i.e., bovine serum albumin (BSA, 66 kDa) and trypan blue (TB, 961 Da) to study the release profile and correlation between different sized model molecules and varying matrix crosslinking density on release behavior.

Further, the reloadability of the system was assessed using fluorescein

isothiocyanate-inulin (FITC-Inulin, 3.9 kDa) and BSA. The system was replenished with 5 ACS Paragon Plus Environment


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drug at regular intervals to provide a constant delivery over an extended period of time. This reloadable system formed of SF was biodegradable, redeeming it of surgical complications for the removal of the implant and also assisting in release of the drugs. Our study aims to develop a scaffold-hydrogel based system that acts a depot for drug delivery. This reservoir was fabricated using silk fibroin scaffolds and silk fibroin hydrogel, creating a blended matrix of the two. The model molecules (BSA, TB and FITC-inulin) were encapsulated in this system and further the molecules release behavior was studied. The reloading ability of our system might allow not only sustainable delivery of drugs but also change of medication after implantation, using different drugs at different stages of the healing process. 2. EXPERIMENTAL SECTION 2.1. Preparation of Aqueous Silk Solution Aqueous solution of silk fibroin (SF) was obtained from Bombyx mori (B. mori) silkworm cocoons according to the procedure described by Mandal et al.44 Briefly, fresh cocoons were collected from Khanapara sericulture farm (Guwahati, Assam, India) and cut into small pieces followed by degumming in 0.02 M sodium carbonate (SRL, India) solution for 20 min to remove sericin, the glue-like component that sticks fibroin chains together. The degummed fiber was then air dried at room temperature and dissolved in 9.3 M LiBr (Sigma, U.S.A.) solution at 60 °C for 4 h followed by dialysis against deionized water for 2 days using 12 kDa membrane (Sigma, U.S.A.) to remove residual LiBr molecules (Figure 1). Final concentration of SF was determined gravimetrically by weighing the remaining solid after drying.


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Figure 1. (A) Schematic illustration of silk fibroin isolation, purification, and matrix fabrication process from B. mori cocoons. (B) Fabrication of model molecule embedded matrices; (i) hydrogel loaded with BSA, (ii) BSA loaded scaffold-hydrogel construct, (iii) hydrogel loaded with trypan blue and (iv) trypan blue loaded scaffold-hydrogel construct. 2.2. Fabrication of Silk Scaffolds, Hydrogels and Scaffold-Hydrogel Hybrid Construct Aqueous derived SF scaffolds were prepared by salt leaching method.31 NaCl salt granules of 800-900 µm were obtained by sieving through metal mesh and added to 8 wt% SF solution at 2:1 (w/v) ratio in cylindrical shaped container. The containers were then left to dry at room temperature and was followed by 12 h ethanol (70% v/v) treatment. Salt leaching was then carried out by immersing them into distilled water for 2 days, with 3-4 changes of water per



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day (Figure 1A). The obtained porous scaffolds (15 mm in diameter and 5 mm in height) were used for in vitro studies or stored in 70% (v/v) ethanol for further studies. To prepare SF hydrogels by sonication method, 5 ml of each 4, 6 and 8 wt% silk solutions were sonicated with a Sonic Vibra-Cell (Sonics & Materials, Inc. U.S.A.), which consisted of the model VC505 power supply, converter and 1/2″ (13 mm) diameter tapered micro-tip. Each SF solution was sonicated for 1 min at 25% amplitude setting and incubated at 37 °C to monitor sol-gel transition visually. BSA and trypan blue impregnated hydrogels were prepared by separately mixing 3 mg/mL BSA (66 kDa) and 5 mg/mL trypan blue (961 Da) (Sigma, U.S.A.) respectively to silk solution prior to sonication (Figure 1B). 2 mL of these sonicated samples were poured into 24 well tissue culture plates and kept undisturbed overnight. The hydrogel blocks (15x5 mm) were then removed from the wells and placed in a closed chamber saturated with ethanol vapor. This condition was maintained for 30 min at room temperature (RT, 25 °C) to ensure greater β sheet induction.16,45 To prepared scaffold-hydrogel hybrid constructs, desired wt% of sonicated SF solution (4, 6 and 8 wt%) was injected into 8 wt% SF scaffolds prior to its gelation The hybrid constructs were kept in room temperature (RT, 25 °C) allowing them to form gel inside. Further the constructs were treated with alcohol vapors for 12 h at RT Similar to biomolecules impregnated hydrogel preparation; BSA and trypan blue of same concentrations (3 mg/mL BSA and 5 mg/mL trypan blue) were injected with hydrogel to fabricate each type of scaffold-hydrogel hybrid construct i.e., scaffold-hydrogel (4 wt%), scaffold-hydrogel (6 wt%) and scaffold-hydrogel (8 wt%) for model molecule release study. 2.3. Gelation Pattern Optimization In order to optimize gelation pattern, 3 mL SF solution of varying concentrations (4, 6 and 8 wt %) was sonicated at 25% amplitude for 30, 45 and 60 s. The phase transition time 8 ACS Paragon Plus Environment

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including translucent time (TT) (refers to the time of change in transparency of solution), viscous time (VT) (time taken to attain the viscous nature of solution) and gelation time (GT) (refers to the time required for sol-gel transformation) were monitored by tube inversion method during the process of hydrogel formation. A total set of 6 (n=6) samples per group were analyzed for this study. 2.4. X-Ray Diffraction (XRD) X-Ray diffraction spectroscopy (D2 Phaser, Bruker, U.S.A.) was carried out to study the extent of induced crystalline structure within B. mori SF scaffold, hydrogel and scaffoldhydrogel hybrid constructs. Samples were analyzed under working condition of Cu Kα radiation (30 kV voltage, 40 mA current having Ni filter) with a scanning rate of 0.5° min-1. The data was analyzed in 2θ angle ranging from 5 to 50° with LYNEXEYE detector system. A total of 3 (n=3) samples per group was analyzed. 2.5. Scanning Electron Microscopy (SEM) Scaffolds, hydrogel and scaffold-hydrogel surface imaging were performed using scanning electron microscopy (SEM) (JSM-5800, JEOL, Japan) with an operating voltage of 20 kV. Cross-sections of the freeze dried samples were prepared by freeze fracture followed by gold sputtering. Surface morphology, pore size and pore interconnectivity were analyzed using SEM. The pore size was determined using Image J 1.40 software by selecting random pores (average of 25 pores). A total of 3 samples (n=3) per group was analyzed. 2.6. Fourier Transform Infrared Spectroscopy (FTIR) To understand the secondary structure in SF scaffold and different wt% of hydrogels (4, 6 and 8 wt%), FTIR analysis was carried out using an infrared spectrophotometer (Nicolet iS 10, Thermo Scientific, U.S.A.) equipped with a thermo-electrically (TE)-cooled deuterated



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triglycine sulfate (DTGS) detector with KBr (potassium bromide) window. To avoid moisture effects, samples were desiccated overnight and prepared KBr disc in the ratio of 9:1(KBr to sample). Spectra were recorded in wavenumber range of 1800-1200 cm-1 by collecting 32 scans with a resolution of 4 cm-1. 2.7. In Vitro Enzymatic Degradation In vitro enzymatic degradation of silk scaffolds, hydrogel and scaffold-hydrogel hybrid constructs was studied by immersing each sample in 2 mL solution of Protease XIV (Sigma, U.S.A.) kept in a 24-well plate placed at 37 °C. Protease solution of 2 U mL-1 activity was freshly prepared in PBS (pH 7.4) with 0.05% sodium azide (Sigma, U.S.A.) to inhibit microbial growth. Similar combinations of SF matrices without enzymes were used as control. The enzymatic solution was replenished with fresh protease solution after every 72 h. For measuring the dry weight, the enzymatic solution was removed and samples were kept overnight at 60 °C in hot air oven. The dry weight was recorded at 0, 7, 14, 21 and 28 days and the percentage degradation of SF matrices was determined by using the formula39 % degradation = [(Wi – Wt )/ Wi] × 100

(1)

Where, Wi is the initial dry weight of SF matrices and Wt is the final weight after 0, 7, 14, 21 and 28 days of enzymatic incubation. The experiment was performed in triplicates (n=3) for each sample group. 2.8. Swelling Behavior of Matrices Swelling behavior of silk scaffolds, hydrogel and scaffold-hydrogel hybrid constructs was analyzed by gravimetric procedure using electronic balance (BSA 224S-CW, Sartorius, Germany). The initial weight (Wi) of silk matrices was measured followed by immersion in phosphate buffer saline (PBS, pH 7.4) at 37 °C for prolonged time (20 h). At predetermined


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time points, the samples were recovered from PBS and the rehydrated weight (Wr) was recorded after wiping of excess PBS from surface area. A total of 6 samples (n=6) per group was analyzed. The swelling ratio of samples was calculated using the following formula39 Swelling ratio = (Wr– Wi)/ Wi

(2)

Where, Wr and Wi indicate rehydrated sample weight and initial sample weight respectively. 2.8.1. Swelling Kinetics Swelling kinetics of SF scaffolds, hydrogel and scaffold-hydrogel hybrid constructs was determined by using the following equation19 f = ktn

(3)

Where, ‘f’ is the fractional water content, ‘k’ is a constant related to SF matrices, ‘t’ is the swelling time and ‘n’ denotes a number which indicates if swelling is due to diffusion or relaxation control. Fractional water content (f) is the ratio of weight of water (wt) at time ‘t’ to the weight of water after attaining swelling equilibrium (w∞). This equation has limitation, as it is only valid for fractional water uptake of 0.6. A total of 6 samples (n=6) per group were analyzed. 2.9. Mechanical Testing Compressive tests of SF scaffolds, hydrogel and scaffold-hydrogel hybrid constructs were performed on Bose Electroforce® 3200 (Bose Corp., U.S.A.) equipped with a 225 N load cell. All the test samples were hydrated at least 30 min before experiments that were carried out in 0.1 M PBS at 37 °C. To test compressive properties, all type of samples were prepared in cylindrical shape measuring 10 mm in diameter and 10 mm in height, according to a modification based on ASTM method F451-95. The compression test was performed



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conventionally as an open-sided method and with a crosshead speed of 0.5 mm m-1. The compressive stress and strain curve were plotted based on the sample height and measured cross-sectional area (nominal 5 mm, measured at 0.01 N tare load), respectively. The compressive modulus was calculated based on a linear regression fitting of a small strain section (2.5-6%) that precedes an identifiable plateau region in all tests. 2.10. In Vitro Release Behavior In vitro release study of model molecules (Bovine serum albumin, BSA and trypan blue, TB) from SF scaffolds, hydrogel and scaffold-hydrogel hybrid constructs was carried out in PBS (pH 7.4) kept at 37 °C. PBS immersed scaffolds, hydrogel and scaffold-hydrogel constructs without model molecule, served as control. The samples were immersed in 15 mL of PBS containing 0.05% sodium azide in a 50 mL centrifuge tube. The entire setup was kept in shaker incubator with a rate of 25 rpm at 37 °C. At different time points (0, 5, 10, 15, 20, 25 and 28 days), 20 µL releasate was collected from each sample and analyzed. The cumulative release of BSA and TB was determined by mixing of 20 µL BSA releasate with 250 µL Bradford’s reagent (Sigma, USA), and 20 µL trypan blue releasate with 180 µL PBS separately in a 96 well plate and recording the absorbance at 595 nm wavelength19 using Tecan infinite M 200 reader with i-control 1.5 software. A total of 6 samples (n=6) per group was analyzed. 2.10.1. Ritger and Peppas Equation to Assess Release Behavior from SF Matrices In order to understand the factors controlling sustained release of model molecules, the absorbance data obtained from sample releasate was fitted to Ritger and Peppas equation: Mt / M∞ = ktn

(4)


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Where, Mt represents the releasate amount at time ‘t’, ‘M∞’ is the releasate amount at time approaching infinity, ‘k’ denotes kinetic constant, ‘t’ is the release time and ‘n’ the diffusion factor which entail about the release mechanism. The above equation can also be correlated with absorbance of a sample releasate because the absorbance, concentration and releasate amount of sample are directly related with each other. 2.11. In Vitro Reloadable Model Molecule Release Study BSA (66 kDa) and FITC-Inulin (3.9 kDa) (each at a concentration of 5 mg mL-1 in PBS, pH 7.4) were used as model molecules to study biomolecule release from the scaffold-hydrogel hybrid constructs. 350 µL of each model molecules was separately loaded into corresponding scaffold-hydrogel hybrid constructs. In brief, desired amount of model molecule was poured into a 1 ml insulin syringe having a needle of 29G (0.33 mm x 12.7 mm). The needle was carefully inserted in to the central core of the construct and model molecule was slowly injected to avoid any air bubbles. This ensured even distribution of molecules within construct. The model molecule loading was repeated thrice upon reaching a plateau region of the previous loading. The model molecules loaded hybrid constructs were placed in 15 mL PBS (pH 7.4) at 37 °C under stirring condition (100 rpm). The study was conducted under air tight conditions to minimize evaporation-mediated PBS loss. A volume of 200 µL was taken for measurement at regular time intervals and replaced with the same volume of PBS as fresh medium to maintain a steady state conditions. Fluorescence measurement at 490 nm and Bradford assay were employed to determine the release of FITC-Inulin and BSA respectively at different time intervals against their corresponding standard curves. The release studies were continued for a time period of 120 h and 240 h for BSA and FITC-Inulin respectively. Dark conditions were maintained for the release study of FITC-inulin throughout to prevent photo-quenching upon exposure to light.



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2.12. Biological Assessment 2.12.1. Isolation and Expansion of Primary Chondrocyte Porcine articular chondrocytes were isolated from 3-4 month-old pigs using modified protocol as described by Akeda et al.46 Cells were obtained from tissue in sequential protease (Sigma, U.S.A.) and collagenase (Sigma, U.S.A.) digestion. Briefly, cartilage was dissected from knee joint surface, finely chopped and washed with PBS (pH 7.4), under sterile conditions. Chopped cartilage tissue was treated with 0.2% (w/v) protease XIV in 5 mL of high glucose Dulbecco’s Modified Eagle’s Medium (DMEM) containing 5% fetal bovine serum (FBS) for 1 h at 37 °C in 5% CO2 incubator followed by 0.018% (w/v) collagenase digestion for overnight, with periodic agitation. The final digest was then centrifuged at 125 g for 10 min and resuspended in high glucose DMEM containing 10% FBS and supplemented with 4 mM glutamine, 2.5 µg mL-1 fungizone, 100 U mL-1 penicillin, 100 µg mL-1 streptomycin (Gibco, Life Technologies, U.S.A.). To isolate single cell suspension, digest was filtered through a 70 µm cell strainer (BD Bioscences, U.S.A.). Cell number and viability were determined by Calcein-AM/ethidium homodimer staining. For initial proliferation studies, isolated chondrocytes were plated at high density in high glucose DMEM and placed at 37 °C in CO2 incubator with 5% CO2 and 85% humidity. After 3 to 4 days incubation, nonadherent cells were removed leaving the adherent cells to achieve 70-80% confluence followed by subsequent expansion in 1:3 ratios. For in vitro experiments, passage three (P3) chondrocytes were used. 2.12.2. Cell Seeding on Scaffold-Hydrogel Hybrid Constructs To evaluate biocompatibility of scaffold-hydrogel hybrid constructs, primary porcine chondrocytes were seeded on each type of scaffold-hydrogel construct (4, 6 and 8 wt%) and only scaffolds were taken as control group. All types of constructs were cut into small pieces 14 ACS Paragon Plus Environment

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(6 mm in diameter and 2 mm in thickness) and sterilized with 70% ethanol followed by vigorous washing in sterile PBS (pH 7.4). Before cell seeding, constructs were conditioned in complete high glucose DMEM for overnight. To monitor cellular behavior, 5 µL aliquots containing 5 x 105 cells were seeded onto each construct. The cell seeded constructs were then left for 3-4 h for initial cell attachment followed by transfer in fresh 12 well plate containing medium. The seeded constructs were incubated at 37 °C with 5% CO2. Media was changed every 3 days over four-week period. 2.12.3. Alamar Blue Cell Proliferation Assay Cell proliferation on SF scaffold and scaffold-hydrogel matrices was monitored using the Alamar blue dye reduction assay (AbD Serotec, U.K.) at specific time period of 1, 7, 14, 21 and 28 days. Primary chondrocytes were seeded at a density of 50,000/matrix kept in a well of a 24-well plate containing DMEM media. The seeded constructs were incubated at 37 °C with 5% CO2 for cell growth and proliferation. At different time intervals absorbance was measured at 570 and 600 nm using a multiplate reader and cell proliferation was calculated according to manufacturer’s protocol. A total of 4 samples (n=4) per group was analyzed. 2.12.4. Live/Dead Assay in Scaffold-Hydrogel Hybrid Constructs After two weeks of culture, cell viability in scaffold-hydrogel hybrid constructs was screened by using live/dead assay kit (Invitrogen, Life Technologies, U.S.A.). Following manufacture’s protocol, constructs were incubated with the live/dead solution at 37 °C for 20 min. The solution is a mixture of three components: 2 mM ethidium homodimer-1 (EthD-1), PBS, and 4 mM calcein-AM. Virtually non-fluorescent cell-permeant calcein-AM enters into cells and consequently is enzymatically converted to intensely fluorescent calcein which is retained inside the live cells producing green fluorescence (ex/em ~495 nm/~515 nm). EthD1 enters into cells through damaged membrane and binds to nucleus generating a bright red 15 ACS Paragon Plus Environment


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fluorescence (ex/em ~495 nm/~635 nm) in dead cells. Live cells, with intact plasma membrane exclude EthD-1. The constructs were washed with sterile PBS (pH 7.4) before imaging. To visualize cellular arrangement, nuclei were stained with Hoechst 33342 (Invitrogen, Life Technologies, U.S.A.) for 30 min under dark condition followed by imaging under fluorescent microscope (EVOS FL cell Imaging System, Life Technologies, U.S.A.). 2.13. Statistical Analysis All quantitative experiments were performed atleast in triplicate and results were plotted as mean ± standard deviation. One-way analysis of variance (ANOVA) with Bonferroni's test was applied for statistical analysis of data. Statistically significant level among groups was accepted at p < 0.05 and in figures, it was indicated as *p ≤ 0.05 and **p ≤ 0.01. 3. RESULTS 3.1. Preparation of Scaffolds, Hydrogel and Scaffold-Hydrogel Hybrid Constructs Scaffold fabrication by NaCl salt leaching method is one of the easiest and effective methods accepted in various fields of tissue engineering. Salt-leached scaffolds were prepared by pouring salt granules of 800-900 µm into a cylindrical vessel containing 8 wt% SF solution. Scaffolds were obtained after leaching of salt granules from SF-salt composites followed by vigorous washing in distilled water and lyophilization. The fabricated scaffolds were porous with interconnected pores that helped in impregnation of hydrogel. Varied concentration of hydrogels (4, 6 and 8 wt%) were prepared by sonication at 25% amplitude for 1 min. Scaffold-hydrogel hybrid constructs were fabricated by injecting sonicated SF solution (in pre-gel condition) into 8 wt% scaffolds. These prepared hydrogel and scaffold-hydrogel hybrid constructs were further used to incorporate model molecules (BSA and TB or FITCinulin in reloadable drug delivery system) in drug delivery assay (Figure 1B).


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3.2. Gelation Pattern Optimization To optimize SF gelation profile after sonication, phase transition time including translucent time (TT), viscous time (VT) and gelation time (GT) were monitored during gelation process (Table 1). Table1. Summary of Phase Transition Time for Sonicated Silk Samples (n=6)

Sonication Time (s) 30 SF concentration (wt%)

45

60

Average phase transition time (min) TT

VT

GT

TT

VT

GT

TT

VT

GT

4

5-10

101

126

1-5

43

57

1

15

24

6

5-10

184

240

1-5

54

68

1

15

21

8

5-10

289

328

1-5

67

100

1

11

17

For 30 and 45 s sonicated SF sample, the mean VT and GT were observed to increase (longer time) with increase of SF concentration (4˂6˂8 wt %), whereas, for 60 s sonicated sample the mean VT and GT showed decreasing (less time) trend with increase in SF concentration (4-8 wt %) (Figure 2C).



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Figure 2. Gelation pattern optimization: the effect of sonication time on different concentration of SF solutions. Plots depicting sonication time of (A) 30 s, (B) 45 s and (C) 60 s. However, with increase in sonication time (30˂45˂60 s) faster gelation time (GT) was observed for SF (4˂6˂8 wt %). For 30 s sonicated sample, the observed mean GT was 120, 240 and 330 min for 4, 6 and 8 wt% SF solution, respectively (Figure 2A). For 45 s sonicated samples the observed mean GT reduced to 60, 70, and 100 min for 4, 6, and 8 wt% SF, respectively (Figure 2B). In case of 60 s sonicated sample, mean observed for GT was


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further reduced to 24, 21 and 16 min for 4, 6 and 8 wt% SF, respectively (Figure 2C). However, in all cases the SF solution lost translucency faster. 3.3. Analysis of Secondary Structure Transition Degree of crystallinity and secondary structure transition within scaffolds, hydrogels and scaffold-hydrogel hybrid constructs were evaluated by XRD and FTIR measurements. In case of scaffolds, hydrogel and scaffold-hydrogel hybrid constructs, two X-ray diffraction peaks were observed at 2θ =21° (major peak) and 24° (minor peak) with a shoulder peak at 41° (Figure 3A). In FTIR analysis, for all SF matrices it was found that vibrational peaks lie in the wavenumber range of 1630-1650, 1520- 1540 and 1220-1270 cm-1 respectively (Figure 3C). 3.4. Scanning Electron Microscopy (SEM) Analysis Matrix morphology is an important criterion for its applicability in drug delivery applications. In this regard, surface morphology, pore size and pore distribution within scaffolds, hydrogel and scaffold-hydrogel hybrid constructs were analyzed using SEM (Figure 3B). Empty SF scaffolds, hydrogels alone and scaffold-hydrogel constructs showed rough surfaces under SEM. Particularly, 8 wt% empty SF scaffolds showed bigger pores throughout with wellconnected inner pores of rounded morphology (Figure 3B, iv). Image J software analysis revealed pore diameter in the range of 700-800 µm as major pores with 100-200 µm smaller interconnected pores. Comparative morphological analysis of 4, 6 and 8 wt% hydrogels showed rough surface along with numerous smaller pores (1-10 µm) distributed throughout. Particularly in case of 4 wt% hydrogel, greater porous architecture was observed compared to 6 and 8 wt% hydrogels (Figure 3B, i-iii). SEM images of all scaffold–hydrogel embedded matrices clearly showed the injected hydrogel occupying available porous spaces between SF



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scaffolds as well as outer surfaces thereby providing rough morphological appearance to the matrices (Figure 3B, v-vii). 3.5. In Vitro Enzymatic Degradation Ability of the fabricated carrier vehicle to degrade overtime in vivo after release of cargo is an important parameter in drug delivery for its safe removal. Similar physiological condition was mimicked in vitro to access SF constructs degradation. In vitro degradation of blank scaffold and scaffold-hydrogel hybrid constructs were studied at 37 °C for 28 days in PBS at physiological pH 7.4 containing protease enzymes (Figure 3D).


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Figure 3. (A) X-ray diffraction spectra of (i) scaffolds and hydrogels only and (ii) scaffoldhydrogel hybrid constructs; (B) SEM images of various SF matrices showing (i) 4 wt% hydrogel, (ii) 6 wt% hydrogel, (iii) 8 wt% hydrogel, (iv) Empty scaffold, (v) scaffoldhydrogel (4 wt%), (vi) scaffold-hydrogel (6 wt%) and (vii) scaffold-hydrogel (8 wt%) hybrid constructs. Scale bar = 500 µm; (C) FTIR spectra of various silk fibroin matrices (scaffold



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and hydrogel only); (D) Degradation plot for SF scaffolds and scaffold-hydrogel hybrid constructs (**p ≤ 0.01). Similar combinations of SF matrices without enzymes were used as controls. Compared to blank scaffolds, scaffold-hydrogel matrices exhibit lower controlled enzymatic degradation. At the end of 28 days, blank scaffold showed the maximum mass loss of ~80%, whereas mass loss of approximate 73%, 45% and 30% was observed in case of scaffold-hydrogel hybrid contracts of 4 6, and 8 wt% respectively. Further, negligible mass loss was observed for all control samples kept in PBS (without enzymes) under similar environmental condition (p ≤ 0.01). 3.6. Swelling Behavior SF Matrices To understand swelling behavior of scaffold, hydrogels and scaffold-hydrogel hybrid constructs were analyzed till attainment of swelling equilibrium in buffered saline at physiological pH (Figure 4).


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Figure 4. (A) Swelling characteristic of various SF matrices in PBS (i) Scaffold only and scaffold-hydrogels hybrid constructs and (ii) hydrogels only; (B) Representative graph demonstrate nature of buffer entry and diffusion controlled swelling of SF scaffold, hydrogels and scaffold-hydrogel hybrid constructs; a=Blank scaffold, b=Hydrogel (8 wt%), c=Hydrogel (6 wt%), d=Hydrogel (4 wt%), e=Scaffold-hydrogel (8 wt%), f=Scaffold-hydrogel (6 wt%) and g=Scaffold-hydrogel (4 wt%) (**p ≤ 0.01). Empty scaffolds alone exhibited faster swelling characteristic with quick attainment of swelling equilibrium within ≈7 h compared to scaffold-hydrogel matrices requiring 35 h with low swelling ratio (p ≤ 0.01) (Figure 4A, i). On the other hand, hydrogels alone showed highest swelling ratio with faster attainment of equilibrium within ≈3 h (Figure 4A, ii). Hydrogels alone swelled 22-27 times as compared to scaffold-hydrogel matrices and twice in



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comparison to empty scaffold. Further, in comparison between hydrogels not much in swelling ratio was observed. Swelling kinetics of SF scaffold, hydrogel and scaffold-hydrogel constructs were assessed using mathematical modeling as depicted in Figure 4B. An effort was made to determine transport mechanism in swelling behavior i.e., either due to diffusion or relaxation control by taking account of fractional water content and swelling time at equilibrium (Table 2). Table 2. Transport Mechanism of Buffer Solution to Different Concentration of SF Matrices SF matrix composition

n

k

Transport mechanism

Empty scaffold

0.0481

0.0856

Fickian diffusion

Hydrogel (4 wt%)

0.0866

0.0294

Fickian diffusion

Hydrogel (6 wt%)

0.0971

0.0299

Fickian diffusion

Hydrogel (8 wt%)

0.0642

0.0451

Fickian diffusion

Scaffold-hydrogel (4 wt%)

0.105

0.145

Fickian diffusion


0.2108

0.1207

Fickian diffusion


0.3949

0.1152

Fickian diffusion

In all groups i.e., empty scaffold, hydrogel alone and scaffold-hydrogel hybrid constructs depicted Fickian diffusion mode of buffer entry/swelling. 3.7. Mechanical Properties Assessment of mechanical strength of fabricated matrices is important to avoid failure of constructs due to shear forces in physiological conditions. Comparative compressive strength assessment of blank silk scaffolds, hydrogel and scaffold-hydrogel hybrid constructs was


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analyzed at 37 °C in fully hydrated state to mimic in vivo physiological conditions (Figure 5A).

Figure 5. (A) Compressive modulus of various SF matrices and (B) strain-stress curves for different concentration of hydrogels and scaffold-hydrogel hybrid constructs (**p ≤ 0.01) Compressive modulus of only scaffolds showed 49.5 ± 10.71 kPa, whereas the value increased to 257.07 ± 26.1, 609.67 ± 63.78 and 643.95 ± 130.54 kPa for scaffold-hydrogel (4 wt%), scaffold-hydrogel (6 wt%) scaffold-hydrogel (8 wt%) respectively. Nevertheless, it was found that the scaffold-hydrogel hybrid constructs had 6-10 times higher compressive 25 ACS Paragon Plus Environment


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modulus than blank scaffolds (p ≤ 0.01). However, not much of a difference in compressive strength was observed between scaffold-hydrogels and hydrogels alone except for 4 wt% scaffold-hydrogels (257.07 ± 261 kPa) showing ~3.3 times higher strength compared to hydrogels alone (76.6 ± 12.35 kPa) (Figure 5A). Further, compressive modulus was found to be directly proportional to protein concentration for both hydrogels alone and scaffoldhydrogels hybrid constructs. Stress strains curves for different concentrations of hydrogels and scaffold-hydrogels hybrid constructs were given in (Figure 5B). Compressive elastic modulus was determined using the regions of stress-strain data between 2.5% to 6%. Improved compressive modulus was observed for scaffold-hydrogel hybrid constructs as compared to hydrogels alone. 3.8. In Vitro Release of Model Molecules In vitro release of model molecules (i.e., BSA and trypan blue) from hydrogels and scaffoldhydrogel hybrid constructs were analyzed in PBS at physiological pH 7.4 as shown in Figure 6. The influential role of impregnated model molecules size, matrix cross-link density (as in various wt% hydrogels), and dual barrier effect (hydrogel-scaffold composite) on release behavior was analyzed from releasate based on absorbance. Sustained release of molecules was observed for initial 4-7 days from all SF samples before reaching equilibrium. A comparative 28 day analysis of BSA and TB release revealed faster release rate of TB than BSA. TB reached to equilibrium within 1-2 days compared to 5-7 days taken by BSA depending on sample types (Figure 6A, i and 6B, i). A clear trend was observed (4 wt% > 6 wt% > 8 wt%) in case of both hydrogels alone and scaffold-hydrogels where release rate was faster in low SF concentration and sequentially decreased in both cases (p ≤ 0.01). This trend is true for both TB and BSA release from all types of SF matrices. However, release rate from 4 wt% SF matrices were much faster compared to 6 to 8 wt% with minimum difference between 6 and 8 wt% when compared in both SF matrix Types. The release behavior of 26 ACS Paragon Plus Environment

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model molecules (Fickian diffusion/degradation or both) from SF matrices was further evaluated by mathematical modeling and fitting the obtained release data into Ritger and Peppas equation (Figure S1, i-ii). In all SF groups i.e., hydrogel alone and scaffold-hydrogel hybrid constructs showed Fickian diffusion mode of drug release behavior (Table 3). Table 3: Transport mechanism for incorporated model molecules from various SF matrices SF matrix composition

BSA

Trypan blue

n

R2

n

R2

Hydrogel (4 wt%)

0.08706

0.844

0.13233

0.7503

Hydrogel (6 wt%)

0.05904

0.9141

0.16468

0.7915

Hydrogel (8 wt%)

0.0411

0.8143

0.16461

0.8173


0.07725

0.9223

0.16836

0.7274


0.03406

0.7896

0.16343

0.7123


0.03102

0.7725

0.17157

0.7491

3.9. In Vitro Release of Model Molecules Using Scaffold-Hydrogel Hybrid Constructs as Reloadable Device Maintenance of structural integrity in physiological condition is prerequisite for application of reloadable drug-delivery device. Scaffolds or hydrogels alone may not be appropriate for sustainable drug delivery systems. Scaffolds tend to possess larger pore size while hydrogels often lose their structural integrity in the physiological medium. In this context, we report our endeavors to fabricate a hydrogel incorporated scaffold to circumvent the afore-stated issues. This system may be envisioned to act as a befitting reservoir for sustained delivery of biomolecular cargo. In this proposed model, hydrogels containing varied silk concentration (4, 6 and 8 wt%) were incorporated in 8 wt% silk scaffolds. The effect of increasing silk 27 ACS Paragon Plus Environment


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concentration on the release behavior of two model biomolecules namely BSA and FITCInulin was assessed over a period of 120 h and 240 h, respectively (Figure 6C). In the first loading, the cumulative % release of BSA for scaffold-hydrogel 4 wt%, 6 wt% and 8 wt% was found to be ̴ 30 ± 3.24%, ̴ 27 ± 2.63% and ̴ 26 ± 3.85% after 3 h and ̴ 55 ± 3.34%, ̴ 38 ± 4.13% and ̴ 36 ± 3.05% after 40 h respectively (Figure 6C, i). Following second loading, the release profile showed a similar behavior attaining plateau at 54 ± 3.41% (4 wt%), 42 ±3.35% (6 wt%) and 24 ± 3.60% (8 wt%) after 40 h. Likewise the release profile after third loading plateaued at ̴ 52 ± 2.12% (4 wt%), ̴ 40 ± 1.13% (6 wt%) and ̴ 20 ± 2.05% (8 wt%) after 40 h. The % release behavior of FITC-Inulin was observed to be similar to BSA, i.e., increasing % of silk hydrogel decreases the % release of model molecules (Figure 6C, ii).


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Figure 6. Effect of silk hydrogel concentration in scaffold-hydrogel hybrid constructs on release behavior of model molecules in phosphate buffer (pH 7.4); (A) BSA, (B) Trypan Blue and (C) reloadable behavior of hybrid constructs, release profile of (C, i) BSA and (C, ii) FITC-inulin (**p ≤ 0.01 and *p ≤ 0.05). After first loading the % release was 62 ± 3.08% (4 wt%), 35 ± 2.63% (6 wt%) and 25 ± 1.93% (8 wt%) after 3 h reaching a plateau at 95 ± 2.18% (4 wt%), 76 ± 1.43% (6 wt%) and 54 ± 1.85% (8 wt%) after 60 h. A very similar release pattern was observed after the second and third loading in each scaffold-hydrogel construct. 3.10. Cell Proliferation Alamar blue assay was carried out to assess primary chondrocyte proliferation on blank scaffold and scaffold-hydrogel matrices. Blank scaffolds (8 wt%) showed highest cell proliferation capacity (4 times) over a period of 4 weeks. On comparing scaffold-hydrogel matrices, 8 wt% scaffolds embedded with 4 wt% hydrogel showed maximum proliferation (~3.7 times) followed by 6 wt% (~3.2 times) and 8 wt% (~3 times) as shown in the figure (Figure 7A).

Further statistically, cell proliferation within 4 wt% hydrogel embedded

scaffold was comparable to blank scaffolds (p ≤ 0.05). 3.11. Cell Morphology and Viability The scaffold-hydrogel hybrid constructs supported the growth and proliferation of porcine chondrocytes over 2 weeks (Figure 7B). Chondrocytes remained attached and maintained their morphology throughout the culture. All three scaffold-hydrogel hybrid constructs (4, 6 and 8 wt%) showed homogeneous cell distribution after 2 weeks. Based on Live/dead assay, approximately 95% cells of the entire cell population were shown to be viable (green color) in all scaffold-hydrogel hybrid constructs.



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Figure 7. (A) Alamar blue assay representing porcine chondrocyte proliferation on SF matrices; (B) Live/dead staining of seeded chondrocytes in scaffold-hydrogel constructs after 2 weeks. (i) scaffold-hydrogel (4 wt%), (ii) scaffold-hydrogel (6 wt%), and (iii) scaffoldhydrogel (8 wt%). Hoechst 33342 staining for nuclei (contrast blue; iv, v, and vi). Scale bar=400 µm. (*p ≤ 0.05). 4. DISCUSSION In contrast to conventional single loading drug delivery systems, the primary aim of the current study was to develop a reloadable drug depot system for sustained delivery applications. Silk fibroin was the material of choice due to its widely reported applications in drug delivery and regenerative medicine in the last few decades.9,11,16,19,24 To achieve sustained and reloadable ability, in most cases the nature of drug molecules or interaction with delivery material is the prime governing factor.47 For reloadable applications, designed reservoir should have enough stability in physiological conditions to maintain its structural 30 ACS Paragon Plus Environment

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integrity so that drugs may be loaded repeatedly in the same cargo. However, there are limitations of only scaffolds as they possess larger pore size that finally may not follow the sustainable pattern in drug delivery rather burst release. On other hand, it is difficult to maintain the structural integrity of hydrogel alone in physiological conditions. To overcome such aforementioned limitations, SF hydrogel was incorporated into the scaffolds that produce a hybrid construct into which model molecules were loaded. Although there are several methods like gas foaming and freeze-drying to make porous scaffolds, salt leaching is one of the easiest and effective procedures that produce highly porous and interconnected pores with tunable mechanical properties.28 Similarly, the porosity and pore size can also be controlled by the size of salt granules applied. In salt leaching procedure, salt granules are partially dissolved in aqueous medium retaining most of the salt part in central portion due to super saturation of silk fibroin solution. As silk fibroin is protein in nature with alternative hydrophobic and hydrophilic domains, its solubilization decreases with the increasing salt concentration. This phenomenon is known as ‘salting out’ effect.48 In presence of high salt concentration, water molecules surrounding protein molecules are removed (termed as water structure maker) that leads to greater interaction between two hydrophobic domains of silk fibroin. This rearrangement of hydrophobic-hydrophilic domains helps in formation of more stable β-sheet structure.49 There are several external inducers (viz., sonication, pH or temperature) that are applied to fabricate silk hydrogel.25,50,51 Gelation time of silk fibroin is considerably very high (days to weeks) without applying any of those external inducers. For instance, lowering the pH or increasing the temperature, gelation time may be reduced to few hours.25,50,51 Sonication is one of those external inducers that can induce gelation within a very short period of time. It is reported that high mechanical vibration produced by sonication causes random formation and collapse of air bubbles. This phenomenon is termed as cavitation. Due to this, local media 31 ACS Paragon Plus Environment


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experiences high temperature (10,000 K), high strain rate (107 s-1), high pressure (200 bar).52,53 All these physical factors cause changes in hydrophobic hydration leading to alteration in hydrophobic-hydrophilic interactions and finally lead to physical cross-linking among the chains (β-sheet formation). However, sonication induced gelation time depends on several parameters like sonication amplitude, pH, presence of ions (K+, Ca++) or concentration and volume of silk fibroin.21 Gelation pattern was studied to ascertain the exact time taken by the gelation processes of various concentrations of silk solution, sonicated for different time intervals. Elucidation of the gelation time aided in successfully injecting the sonicated samples into the porous scaffold prior to gelling, to form the desired scaffold-hydrogel hybrid constructs. The sonication of silk solution causes a series of sequential changes before turning into a hydrogel. These stages are translucent stage, viscous stage and gelation stage and the time taken are called translucent time (TT), viscous time (VT) and gelation time (GT) respectively. The results obtained from the experiment indicated the inverse relationship between sonication time and mean gelation time. It can be elaborated as the increased sonication time causes decrease in mean GT which has been reported previously.21 Results of 30 s and 45 s sonicated samples also revealed direct proportionality between SF concentration and GT (Figure 2A and 2B). The outcome of this result was due to the high concentration of SF solution that could resist the propagation of sonication wave. This may be stated differently that sonication time was not enough to induce gelation in SF solutions of higher concentrations.21 In contrast, the retrospective decrease in GT for 60 s sonicated sample indicated the correlation between increased sonication time and SF concentration (Figure 2C). One of the reasons for reduced GT of 8 wt%, 60 s sonicated samples may be due to the faster rate of β-sheets induction as such solutions have more number of SF molecules than 4 and 6 wt%, and 60 s sonication provides enough energy to expose more 32 ACS Paragon Plus Environment

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number of hydrophobic surfaces in 8 wt% silk solutions. However, after sonication, varying concentrations of SF solutions (in pre-gel condition) were injected into 8 wt% SF scaffolds to make scaffold-hydrogel hybrid constructs. These different concentrations of hydrogel (4, 6 and 8 wt%) were used for tunable release of model molecules from the hybrid matrices. The concentration of SF molecules influences the crosslinking density and consequently the pore size and porosity, thereby offering a route to tune the release profile of model molecules. The obtained X-ray diffraction peaks indicated secondary structure transition in SF matrices which were consistent with previously reported result.9,54 Another study indicated that peak intensity at 23° and 26° correspond to 50% of β-sheet structure transition.55 The obtained Xray diffractogram showed that the diffraction peak intensity increased with increase in SF concentration (Figure 3A). Further, SF secondary structure transition was determined by analysis of amide vibration peaks in the wavenumber range of 1200-1700 cm-1. SF polypeptide chain is composed of various amide bonds which play important role in forming cross-linking structure within SF matrices. For pure SF, the absorption band resides in the range of 1640, 1545 and 1265 cm-1 for amide I (C=O stretching), amide II (N-H deformation and C-N stretching) and amide III (C-N stretching and N-H deformation) respectively.9 In this study, bands were obtained in the wavenumber range of 1630-1650 and 1520- 1540 cm-1 which was also consistent with previously reported study (Figure 3C).28,54,56,57 The bands at 1630-1650 and 1520-1540 cm- 1 corresponded to the silk II structure transition with some random structure. The obtained FTIR spectral bands also revealed the relationship between SF concentration and spectral peak height; it represents the peak height increases with increasing silk concentration. These obtained results corresponded to silk secondary structure transition along with some random coil conformation. Thus, the secondary structure in SF scaffolds, hydrogel and scaffold-hydrogel hybrid constructs was confirmed by XRD and FTIR spectra. 33 ACS Paragon Plus Environment


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Surface characterizations in terms of pore size or interconnectivity of SF scaffolds depend on fabrication procedures (e.g., particulate leaching or freeze drying techniques) and protein concentration.9,58-60 For instance, it has been reported that pore size decreases as the SF concentration increases in scaffold when prepared at a particular pre-freezing temperature.9 However, for salt leaching procedure, pore size depends on the size of salt granules used. Usually pore size was observed to be smaller than the size of salt granules applied. In this case, the 8 wt% salt leached scaffolds (serving as the control) exhibited pore size of about 700-800 µm where the used salt granules size was 800-900 µm (Figure 3B, iv). The reduced pore size was mainly due to dissolution of NaCl particles surface in aqueous SF solution.58 In case of SEM images, it was visible that less concentrated SF had lesser crosslinking density and more water content. The obtained result supported the greater porous architecture as observed in freeze-dried 4 wt% hydrogel. Furthermore, 6 and 8 wt% hydrogel contain more SF molecule and so have high cross-link density and less water content thereby provide more rigid appearance to the matrix with smaller pore architecture (Figure 3B, i-iii). Swelling ratio of SF matrices was found to be influenced by SF concentration present in the matrix system. More concentrated hydrogel and scaffold-hydrogel hybrid constructs showed more swelling ratio. Our results were consistent with previously reported findings which reveal that the degree of swelling increased with increase in SF content.61 In scaffoldhydrogel hybrid constructs, the injected hydrogel occupied porous spaces of scaffold and thus, it controlled buffer imbibing rate of the matrices. Comparative analysis of swelling ratio for the scaffold-hydrogel hybrid constructs and the hydrogel alone revealed reduced swelling ratio of the former (Figure 4A). One possible cause for this observation could be the presence of extra wrapping and robust cage like spaces within the scaffold for injected hydrogel. Thus the regulated entry and uptake of PBS buffer by matrices allowed limited expansion and swelling capacity of scaffold-hydrogel hybrid constructs. Increased swelling 34 ACS Paragon Plus Environment

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ratio for the 8 wt% SF hydrogels may be due to less water content within more concentrated SF matrices along with their high osmotic potential. Another study revealed that hydrogel can quickly imbibe more than five times of their weight of water in a short duration.62 The nature of water entry into SF matrices was determined by taking natural logarithm of the equation (3). The equation provides a linear plot between ‘ln f’ versus ‘ln t’ with slope (n) and intercept, ‘ln k’ as shown in Figure 4B. For cylindrical SF matrices, the value of n below 0.5 indicates diffusion controlled swelling (Fickian diffusion) nature of SF matrices.19,61 The values of n and k are shown in Table 2, suggestive of diffusion-dictated water absorption and swelling of the matrices. Protease XIV, a biocatalyst cleaves silk fibroin indiscriminately at multiple sites.14,39,63 These enzymes act on biomaterials in two steps, first step involves the adsorption of enzyme on the biomaterial surface and second pertains to bond hydrolysis.64,65 SEM analysis of scaffolds revealed the bigger pores interconnected via numbers of smaller pores which facilitated the easy entry for enzyme and thus its greater accessibility to sites for interaction within the scaffolds. Previous degradation studies have revealed that silk matrices have more β-sheet content with crystal-noncrystal alternate structure and more hydrophilic interaction and undergo faster degradation.56 The XRD and FTIR analysis confirmed the presence of both crystal-noncrystal alternate structures in the SF matrices reported in the present study. Therefore, factors like porous and interconnected morphology along with presence of both crystal and non-crystal structure play important role in faster enzymatic degradation of the scaffold.20,24 Beside these, degree of crosslinking and insolubility of matrix backbone, processing condition of silk, SF concentration and preparation methodology are critical determinants of the rate of enzymatic degradation.64,66 The scaffold-hydrogel hybrid constructs exhibited reduced enzymatic degradation rate due to limited entry of enzyme into the matrix and thereby reducing the surface area of the matrix, exposed to the enzyme. At the 35 ACS Paragon Plus Environment


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end of 28 days, scaffolds showed maximum mass loss of ~80 wt% whereas the values for scaffold-hydrogel hybrid constructs of 4 wt%, 6 wt% and 8 wt% were ~73%, ~45% and ~30% respectively (Figure 3D). The comparatively hindered mass loss of 6 wt% and 8 wt% scaffold-hydrogel hybrid constructs may be due to more cross-link density, less amount of crystal-noncrystal alternate structure and smaller pore size, thereby permitting limited entry of protease enzyme inside the matrix. Further, negligible mass loss was observed for control sample kept in PBS under similar environmental condition. The degradation rate of SF matrices has also been observed to contribute to the controlled release rate of incorporated molecules.67 High mechanical strength is a desirable property of in vivo implantable silk scaffolds, hydrogel and scaffold-hydrogel hybrid constructs used for tissue engineering applications, sustained drug delivery or reloadable purpose, so that these can withstand the mechanical shearing of body part. In order to prepare high mechanical strength matrices along with attributes of control drug release and cytocompatibility, the SF hydrogel was infused into porous scaffolds to form scaffold-hydrogel hybrid constructs. Comparative studies were performed to determine the mechanical strengths of silk scaffolds, hydrogels and scaffoldhydrogel hybrid constructs in their fully hydrated state (Figure 5A). Observations showed that scaffold-hydrogel hybrid constructs had about 6-10 times more compressive modulus compared to blank scaffolds. The results implied the dependency of the compressive modulus on SF concentrations. The mechanical strength was found to increase with the increasing degree of cross-links of the matrices.68 A previous report also revealed the enhanced mechanical property of a quickly formed hydrogel from highly concentrated silk solution.21 More concentrated silk solution formed high cross-linked matrix system with smaller pore, as observed through SEM images (Figure 3B). In case of blank scaffold, bigger pore morphology resulted in uneven load distribution and consequently the reduced compressive 36 ACS Paragon Plus Environment

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modulus of scaffolds. Pore morphology and their distribution play a critical role in determining mechanical strength of the matrices.54,69 Comparative analysis of compressive modulus of blank scaffolds and hydrogel (4 wt%) revealed that 4 wt% hydrogel had higher compressive modulus, attributable to the small pore size with even load distribution in the hydrogel. In case of 6 and 8 wt% hydrogel, they have high cross-linking density with smaller pore architecture which again improved their mechanical properties. The hydrogel infusion into porous scaffold improved the mechanical strength of scaffold-hydrogel hybrid constructs, further supported by swelling behavior analysis of scaffold-hydrogel constructs (Figure 4). It showed that the swelling ratio of scaffold-hydrogel hybrid constructs was far less than the swelling ratio of their hydrogel counterparts. This indicated the hydrogel infusion into scaffold made it a robust matrix system with reduced swelling ability and limited expandability of the scaffold-hydrogel hybrid constructs. Further comparative analysis of hydrogel (8 wt%) and scaffold-hydrogel (8 wt%) hybrid constructs revealed no significant differences of compressive modulus in between them. The outcome of this result may be due to manual incorporation of hydrogel into scaffolds and thereby, uneven distribution followed by sparse hydrogel formation that associated with insignificant changes in compressive modulus. The release of model molecules from hydrogel is governed by several factors including molecular weight and nature of drugs (hydrophobicity or hydrophobicity), interaction of drug molecules with hydrogel, nature of hydrogel (pore size, degree of cross-linking), and solvent type.70-73 Previous reports indicate that with the increase in silk concentration the percentage of ß-sheets increases in response to sonication that further results in decrease in the swelling capability.21 The increased level of β-sheets resulted in a highly cross-linked polymeric network that finally resulted in sustained release of model molecules. BSA and trypan blue were used to study the in vitro release of these model molecules from both hydrogels and 37 ACS Paragon Plus Environment


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scaffold-hydrogel hybrid constructs in PBS buffer (pH 7.4) (Figure 6A and 6B). Initial burst release was observed from all samples due to adherence of model molecules on the surface of matrices. This weak physical interaction attributes to the initial burst release, after this the sustained release phase is noticed. The observed release rate of model molecules from 4 wt% hydrogel was more than 6 and 8 wt% hydrogel. This could be possibly due to comparatively less amount of SF in 4 wt% hydrogel with lesser cross-linking density, bigger pore structures, eventually resulting in enhanced diffusion of the biomolecules. The previous report showed that the cumulative release rate of loaded drug molecules decreased with the increase in SF concentration.74,75 Again, if the hydrogels (4, 6 and 8 wt%) containing model molecules was injected into the scaffolds then, release rate of molecule was further decreased than their hydrogel counterparts. The result clearly indicated the prominent role played by scaffold in controlling the release rate of model molecules. Here, the scaffolds provided extra wrapping and layering around the hydrogel in scaffold-hydrogel hybrid constructs and thus it also improved the mechanical strength besides controlling of matrix degradation rate. Both BSA and TB followed similar release pattern from their respective composition and matrices type. A comparative 28 days releasate analysis of BSA and TB release profile revealed faster release rate of TB than BSA. One of the most probable causes for faster release of TB may be contributed by its smaller molecular size with faster diffusion rate from the matrices despite of similar cross-link density, porosity and pore size of matrices containing BSA. Controlled and sustained release of BSA molecules was observed as attested by Bradford’s assay. In both cases (BSA and TB), similar release pattern was obtained in terms of absorbance. Overall a decreasing order in terms of the release pattern was obtained for hydrogel (4wt %), scaffold-hydrogel (4 wt%), hydrogel (6 wt%), hydrogel (8 wt%), scaffold-hydrogel (6 wt%) and scaffold-hydrogel (8 wt%) respectively. Attributes like reduced swelling ratio, improved mechanical strength, controlled enzymatic degradation rate, sustainably controlled model 38 ACS Paragon Plus Environment

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molecule release rate along with the dual soft and hard nature of the matrices endow special niche to the fabricated biomaterial for drug delivery and tissue engineering applications. The Ritger and Peppas equation model was performed in order to determine the factor majorly responsible for release of incorporated model molecules. For all these SF matrices, the obtained value of ‘n’ lied between 0.03 to 0.17 which was indirectly related to the nature and size of model molecules along with matrix crosslinking density. If value of ‘n’ approaches 0.5, it indicates diffusion type of release mechanism, whereas a value of ‘n’ between 0.5 and 1 is suggestive of a coupled Fickian diffusion and polymer degradation to govern the release behavior.19 Thus from Table 3, we found that the obtained values of ‘n’ are less than 0.5 suggesting diffusion based release of model molecule. This was further supported by the negligible degradation observed for the control sample immersed in PBS buffer (pH 7.4) and swelling behavior analysis from Table 2. It showed the values of ‘n’ lies below 0.5 which suggests the Fickian diffusion type of transport mechanism. However, the same types of SF matrices were further applied in reloadable purposes (Figure 6C). The release profile of BSA molecule showed a burst release at 3 h which was possibly due to the presence of surface adsorbed molecules (Figure 6C, i). However, a complex polymeric network within the hybrid constructs containing higher % of hydrogels led to decrease in the protein release. After every successive reloading of the BSA molecule, the release plateau was attained with similar values indicating an effective reloadable system. However, in case of the 8 wt% hydrogel the maximum release was moderately decreased after each successive loading which may be due to the increased polymeric network of higher concentration protein that interfere in release behavior. The release profile of FITC-Inulin was analogous to that of the BSA (Figure 6C, ii). However, an overall higher cumulative release was observed at each successive loading. This was possibly due to lower molecular weight of FITC-Inulin that facilitated its easy release 39 ACS Paragon Plus Environment


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from the complex polymeric network. Further, the release behavior of these model molecules may be correlated to the swelling capacity of the hydrogels. Scaffold-hydrogel constructs with lower swelling capability are expected to release lower amount of model molecules. The results of the release study were in agreement to the afore-stated statement, where constructs with higher % hydrogel showed less swelling and further less release. The in vitro degradation study further supported the integrity of the hybrid system, forwarding it as potential reservoir for repeated loading of biomolecules/drugs for confined and timely targeted delivery to a region of interest Finally, cell proliferation studies were performed to check the cytocompatibility of these matrices over a period of 4 weeks (Figure 7). Alamar blue assay was suggestive of the influence of matrix-roughness and compositional variation of scaffold-hydrogel hybrid constructs on the initial cell attachment. The maximum cell proliferation was observed in case of blank scaffolds compared to hybrid constructs. This result may be due to the high surface area of blank scaffolds that assisted in initial cell attachment while the porous microenvironment

ensured

nutrient

exchange,

thereby

supporting

maximum

cell

proliferation. However, among the hybrid constructs, scaffold-hydrogel (4 wt%) showed the highest cell proliferation. The higher compactness of 6-8 wt% scaffold-hydrogel constructs may be projected as a possible reason for lower cellular infiltration and consequently lesser cellular growth. 5. CONCLUSION In this work, we have successfully fabricated cytocompatible and biodegradable silk fibroin based hybrid construct as drug-reloadable 3D matrices. The logarithmic release profile of model molecules exclusively showed a diffusion process which was greatly influenced by extra wrapping layer of scaffolds onto the hydrogel and its interaction with the model


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molecules. Further, the controlled release of model molecule for sustained period was affected by pore size and porosity of the SF matrices. Molecule-size played an important role in determining release rates as the smaller molecules were observed to be delivered faster. The fabricated hybrid constructs acted as reloadable depot system for repeated and multiple drug delivery applications. This may serve as a reloadable depot system for drug delivery and tissue engineering applications liberating it of surgical complications for the removal of the implants. ACKNOWLEDGEMENTS The authors greatly acknowledge funding support to BBM from Government of India through Department

of

Biotechnology

(BT/PR6889/GBD/27/490/2012

and

BT/548/NE/U-

Excel/2014), Department of Science and Technology (IFA-13 LSBM-60 and SB/FT/LS213/2012). K. Lakshminath is acknowledged for help in mechanical testing. CIF facility at IIT Guwahati is greatly acknowledged for high end instrument support. Contribution of Mr. Yogendra Pratap Singh and Dr. Rocktotpal Konwarh, IITG is appreciated for their valuable suggestions during manuscript preparation.

SUPPORTING INFORMATION Mathematical model using Ritger and Peppas equation to determine factors for control release of incorporated model molecules including BSA and trypan blue (TB).

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For Table of Contents Only

Reloadable Silk-Hydrogel Hybrid Scaffolds for Sustained and Targeted Delivery of Molecules Saket Kumar Singh1, Bibhas Kumar Bhunia1, Nandana Bhardwaj2, Sween Gilotra1, Biman B. Mandal1* 1

Biomaterial and Tissue Engineering Laboratory, Department of Biosciences and

Bioengineering, Indian Institute of Technology Guwahati, Guwahati – 781 039, India 2

Biological and Chemical Sciences Section, Life Sciences Division, Institute of Advanced

Study in Science and Technology, Guwahati – 781 035, India

*Corresponding author, E-mail: [email protected]; [email protected] Tel: +91-361-258 2225



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Figure 1. (A) Schematic illustration of silk fibroin isolation, purification, and matrix fabrication process from B. mori cocoons. (B) Fabrication of model molecule embedded matrices; (i) hydrogel loaded with BSA, (ii) BSA loaded scaffold-hydrogel construct, (iii) hydrogel loaded with trypan blue and (iv) trypan blue loaded scaffold-hydrogel construct. 149x130mm (200 x 200 DPI)

ACS Paragon Plus Environment

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Figure 2. Gelation pattern optimization: the effect of sonication time on different concentration of SF solutions. Plots depicting sonication time of (A) 30 s, (B) 45 s and (C) 60 s. 150x130mm (300 x 300 DPI)



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Figure 3. (A) X-ray diffraction spectra of (i) scaffolds and hydrogels only and (ii) scaffold-hydrogel hybrid constructs; (B) SEM images of various SF matrices showing (i) 4 wt% hydrogel, (ii) 6 wt% hydrogel, (iii) 8 wt% hydrogel, (iv) Empty scaffold, (v) scaffold-hydrogel (4 wt%), (vi) scaffold-hydrogel (6 wt%) and (vii) scaffold-hydrogel (8 wt%) hybrid constructs. Scale bar = 500 µm; (C) FTIR spectra of various silk fibroin matrices (scaffold and hydrogel only); (D) Degradation plot for SF scaffolds and scaffold-hydrogel hybrid constructs (**p ≤ 0.01). 149x184mm (200 x 200 DPI)


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Figure 4. (A) Swelling characteristic of various SF matrices in PBS (i) Scaffold only and scaffold-hydrogels hybrid constructs and (ii) hydrogels only; (B) Representative graph demonstrate nature of buffer entry and diffusion controlled swelling of SF scaffold, hydrogels and scaffold-hydrogel hybrid constructs; a=Blank scaffold, b=Hydrogel (8 wt%), c=Hydrogel (6 wt%), d=Hydrogel (4 wt%), e=Scaffold-hydrogel (8 wt%), f=Scaffold-hydrogel (6 wt%) and g=Scaffold-hydrogel (4 wt%) (**p ≤ 0.01). 150x129mm (300 x 300 DPI)



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Figure 5. (A) Compressive modulus of various SF matrices and (B) strain-stress curves for different concentration of hydrogels and scaffold-hydrogel hybrid constructs (**p ≤ 0.01). 150x126mm (300 x 300 DPI)


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Figure 6. Effect of silk hydrogel concentration in scaffold-hydrogel hybrid constructs on release behavior of model molecules in phosphate buffer (pH 7.4); (A) BSA, (B) Trypan Blue and (C) reloadable behavior of hybrid constructs, release profile of (C, i) BSA and (C, ii) FITC-inulin (**p ≤ 0.01 and *p ≤ 0.05). 150x116mm (300 x 300 DPI)



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Figure 7. (A) Alamar blue assay representing porcine chondrocyte proliferation on SF matrices; (B) Live/dead staining of seeded chondrocytes in scaffold-hydrogel constructs after 2 weeks. (i) scaffoldhydrogel (4 wt%), (ii) scaffold-hydrogel (6 wt%), and (iii) scaffold-hydrogel (8 wt%). Hoechst 33342 staining for nuclei (contrast blue; iv, v, and vi). Scale bar=400 µm. (*p ≤ 0.05). 149x201mm (200 x 200 DPI)


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Table of content graphic 85x45mm (300 x 300 DPI)


Reloadable Silk-Hydrogel Hybrid Scaffolds for Sustained and

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