Exploring Gelation and Physico-chemical Behavior of In-situ

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Tissue Engineering and Regenerative Medicine

Exploring Gelation and Physico-Chemical Behavior of In-situ Bioresponsive Silk Hydrogels for Disc Degeneration Therapy Bibhas Kumar Bhunia, and Biman B. Mandal ACS Biomater. Sci. Eng., Just Accepted Manuscript • DOI: 10.1021/acsbiomaterials.8b01099 • Publication Date (Web): 23 Dec 2018 Downloaded from http://pubs.acs.org on December 26, 2018

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ACS Biomaterials Science & Engineering

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Exploring Gelation and Physico-chemical Behavior of In-situ Bioresponsive Silk Hydrogels

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for Disc Degeneration Therapy

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Bibhas K. Bhunia and Biman B. Mandal*

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Biomaterial and Tissue Engineering Laboratory, Department of Biosciences and Bioengineering,

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

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Author for correspondence: *[email protected], *[email protected]

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

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Abstract

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Hydrogels have received considerable attention in the field of tissue engineering owing to their

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unique structural and compositional resemblance to the highly hydrated human tissues. In addition,

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controlled fabrication processes benefit them with desirable physicochemical features for

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injectability in minimally invasive manner and cell survival within hydrogels. Formulation of

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biologically active hydrogels with desirable characteristics is one of the prerequisites for

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successful applications like nucleus pulposus (NP) tissue engineering to address disc degeneration.

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To achieve such a benchmark, in this study two naturally derived silk fibroin proteins (Bombyx

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mori, BM SF and Antheraea assamensis, AA SF) were blended together to allow self-assembly

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and transformation to hydrogels in absence of any cross-linker or external stimuli. A

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comprehensive study on sol-gel transition of fabricated hydrogels in physiological fluid

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microenvironment (pH, temperature and ionic strength) was conducted using optical and

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fluorescence analysis. Tunable gelation time (~8-40 min) was achieved depending on

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combinations. The developed hydrogels were validated by extensive physicochemical

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characterizations which include confirmation of secondary structure, surface morphology,

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swelling and degradation. Mechanical behavior of the hydrogels was further analyzed in various

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in vitro-physiological-like conditions with varying pH, ionic strength, diameter, storage time and

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strain values to determine their suitability in native physiological environments. Rheological study,

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cytocompatibility using primary porcine NP cells and ex-vivo biomechanics of hydrogels were

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explored to validate their in situ applicability in minimally invasive manner towards potential disc

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regeneration therapy.

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Keywords: Silk hydrogel, self-assembly, ex-vivo biomechanics, nucleus pulposus, tissue

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engineering 2 ACS Paragon Plus Environment

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1.

Introduction

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Hydrogels consist of three-dimensional networks of hydrophilic polymers that either crosslink

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covalently or physically interact (intra- and inter-molecular attraction) with each other 1. It has the

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ability to absorb large quantity of water or biological fluid and swell up to several hundred times

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without dissolving the polymers. In swollen state, hydrogels become semi-solid and elastic,

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resembling to the living tissues or extracellular matrix surrounding them that attract attention of

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the tissue engineers 2-3.

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Over the past few years, in situ gelation systems have received considerable accolades in the field

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of tissue engineering and regenerative medicine (TERM)

4-5.

In in situ gelation systems, the

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polymers remain in fluid state prior to administration and transform to gel under normal

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physiological conditions that trigger polymer-polymer or polymer-solvent interactions. These

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physiological conditions influence the gelation rate and the final gel strength which are often

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critical to their functions 6. There are several natural and synthetic polymers that are being utilized

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in the recent past for in situ applications in tissue engineering and drug delivery systems. For

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instance, pectin is a natural polymer (polysaccharide) that readily forms gel in presence of calcium

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ions in the stomach following oral administration

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thermally reversible gelation behavior by lateral stacking of chains. It forms gel under

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physiological conditions presenting its potential applications in ocular, intraperitoneal, oral and

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rectal drug delivery 9-11.

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Silk fibroin (SF) is a naturally derived protein polymer that is being used as a biomaterial in various

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biomedical applications

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tunable biodegradability and mechanical properties, and minimal immunogenicity 18. SF can be

12-17

7-8.

Partially degraded xyloglucan shows

due to its easy processability, biocompatibility, hemocompatibility,



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isolated from both mulberry (i.e., Bombyx mori) and non-mulberry (e.g., Antheraea assamensis,

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Antheraea mylitta, Philosamia ricini etc.,) sources

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assembly giving rise to the formation of β-sheets which results in hydrogel formation. Silk

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hydrogel can be fabricated following several methods that include use of chemicals e.g., salts (Ca2+

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and K+ ions), organic solvents (alcohol), surfactants (sodium dodecyl sulphate), large polymeric

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agents (poly-ethylene glycol), low pH and high pressure CO2 or applying physical stimulus, e.g.,

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temperature, shear force, ultrasonication and electric filed 20-24. In a very recent study, Yang et al.,

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reported the efficiency of sodium oleate to induce rapid gelation in SF which might be compatible

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with in situ gelation applications 25. Similarly, Bragg et al., developed an in situ hydrogel platform

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using sonicated or genipin cross-linked SF and gelatin or heparin conjugated gelatin for sustained

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release of growth factors 26. The gelation time of SF solution increases in absence of these non-

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physiological stimuli. External stimuli like sonication or vortexing reduce the gelation time but

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increase the chances of shearing and contamination in vitro. Furthermore, these techniques require

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sophisticated and expensive instruments which again increases the total cost of hydrogel

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fabrication. In the current study, we utilized the self-gelation capability of two different varieties

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of SFs, when blended, as a platform for disc regeneration therapy via in situ gelation.

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Degenerative disc disease (DDD) is often related to low back pain (LBP). Approximately, 80 %

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of the population suffers from LBP in their life time and this issue now-a-days is associated with

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healthcare expenditure up to several billion dollars 27. An intervertebral disc consists of two distinct

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regions; ring like annulus fibrosus (AF) that confines gelatinous nucleus pulposus (NP) in the

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center 28. The NP gel contains ~90 % of water and rest is a composite of protein and polysaccharide

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in the ratio of 1:2 29. NP tissue is not only composed of extracellular matrix (ECM) components,

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but also contains around 4×106 NP cells/ml which is quite low than the surrounding AF tissue

19.

Aqueous solutions of SF undergoes self-


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(~9×106 cells/ml)

30.

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complex structure compared to AF cells 31. Though these cells have low mitotic and regeneration

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capability, they can survive under certain stress conditions like low pH, hypoxia and low glucose

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levels. In disc degeneration, a larger portion of NP cells undergo apoptosis and senescence that

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finally lead to the overall loss of NP matrix followed by decrease in water content 32-33. In a healthy

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disc, NP gel releases pressure and water molecules to the surrounding AF upon compression and

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water molecules in turn get back to the NP upon decompression, driven by osmotic force. This

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pump effect is very crucial to maintain the disc biomechanics (unconfined compressive modulus

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of 3–6 kPa) 34-35. However, this load transfer mechanism gets disrupted in disc degeneration as the

These cells are rounded in morphology with larger cytoplasm and more

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water content of NP gel gradually decreases resulting in declined hydrostatic pressure 36.

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Currently, attempts are being made worldwide by various research groups to develop injectable

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hydrogels using materials such as collagen, alginate, chitosan and hyaluronic acid (HA) as

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scaffolds for NP gel replacement because of their similarity to the native NP tissue 37-40. Most of

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them are used either alone (acellular) or in combination with cells or growth factors in order to

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gelate in the site of NP tissue in situ. For instance, Jalani et al., developed an in situ thermogel

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made of chitosan and HA, crosslinked with genipin and β-glycerophophate that settled for gelation

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within 32 min 41. In a very recent strategy, decellularized NP tissue was being used as in situ gelling

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platform for NP tissue replacement. The decellularized NP tissue reached the gelation point within

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45 min at physiological temperature in presence of fibrillar collagen which was very much similar

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to the native NP tissue 42. However, these in situ hydrogel systems have certain limitations like

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high gelation time, interference of crosslinkers or synthetic counterparts that may pose some

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adverse effects and weak mechanical strength which fails to provide sufficient physical support to

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withstand load in real situation. Moreover, these hydrogel systems display relatively faster 5 ACS Paragon Plus Environment


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degradation rate that cannot be well matched to the slow secretion rate of NP extracellular matrix

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(ECM).

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To address these relevant issues, the present study illustrates the rapid gelling property of aqueous

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solution of SF, formed via blending of two different varieties of silk (i.e., Bombyx mori and

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Antheraea assamensis). The hydropathicity of these two types of SF is different due to their amino

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acid composition, sequence and arrangement 43. The protein chains in aqueous SF solution self-

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assembled and transformed into gel (due to alteration of the hydrophobic-hydrophilic

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microenvironment) in absence of any cross-linker or given external stimulus. The purpose of this

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study was to design a suitable combination of these two SFs for in situ applications in NP

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replacement. The gelation process was optimized under in vitro physiological microenvironment

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(i.e., physiological temperature, pH and ionic strength), thereby implying a special emphasis on

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gelation mechanism on the basis of optical and fluorescence studies. The hydrogels so fabricated

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were extensively characterized in terms of mechanical behavior, rheological properties and ex-vivo

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bio-functional assessments conducted under in vitro physiological conditions. Furthermore,

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cytocompatibility study using primary porcine NP cells validated its in situ applicability in disc

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regeneration therapy.

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2.

Materials and Methods

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2.1.

Preparation of aqueous solution of silk fibroin (SF)

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Aqueous solution of SF from both B. mori (mulberry SF) and A. assamensis (non-mulberry SF)

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were isolated following previous studies

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degummed in 0.02 M sodium carbonate (Na2CO3) followed by dissolution in 9.3 M lithium

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bromide (LiBr) and subsequent dialysis to purify the protein. Contrarily, silk glands of A.

19, 44.

In brief, B. mori cocoons were chopped and


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assamensis were collected and dissolved in 1 % (w/v) sodium dodecyl sulfate (SDS) followed by

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dialysis at 4 °C. The protein concentration from both the sources was measured gravimetrically

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and stored at 4 °C for future use.

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2.2.

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SF isolated from both silk varieties i.e., B. mori and A. assamensis were blended in different ratios

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maintaining the final protein concentration at 3 % (w/v), as summarized in Table 1. The blended

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protein solution was then incubated at different physicochemical conditions (viz., temperature,

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protein concentration, molecular weight of protein and ionic strength) for gelation pattern

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analysis.

Formation of hydrogels

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Table1: Preparation of different ratios of SF blends Sample code

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BM SF g % (w/v)

AA SF g % (w/v)

BM3.0 3.0 AA3.0 0 BM0.5/AA2.5 0.5 BM1.0/AA2.0 1.0 BM1.5/AA1.5 1.5 BM2.0/AA1.0 2.0 BM2.5/AA0.5 2.5 BM = Bombyx mori, AA = Antheraea assamensis

0 3.0 2.5 2.0 1.5 1.0 0.5

Final protein concentration g % (w/v) 3 3 3 3 3 3 3

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2.3. Spectroscopic studies on the silk blends

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2.3.1. Gelation kinetics by visible light spectroscopy

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To study the effect of different parameters on silk blends in hydrogel formation, turbidity change

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during gelation process was assessed by UV-Visible spectroscopic measurements. 200 µl silk 7 ACS Paragon Plus Environment


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blends were taken per well in a 96 well-plate (Tarsons, India) and absorbance at 600 nm were

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measured using a multiplate-reader equipped with temperature control (Tecan Infinite M200,

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Switzerland). The effect of temperature in gelation was monitored at 2 5 °C, 37 °C and 42 °C for

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a period of ~70 min at an interval of 3 min. BM3.0 and AA3.0 (without blending) were used as

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controls. The effect of protein concentration in gelation was studied by blending SF solutions in

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1:1 ratio in an increasing order from 0.5 to 2.0 % w/v for the gelation pattern analysis, while

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keeping one type of SF protein constant (2 %, w/v) and by increasing the other SF protein

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concentration gradually (0.5 to 3.0 %, w/v). The effect of ion concentration and pH was studied

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on 3.0 % (w/v) SF blends (BM3.0:AA3.0, 1:1) at different ion Ca2+, K+ concentrations (1-4 mM,

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adjusted by 1M aqueous stock of KCl and CaCl2); at different pH (7.4 and 9.0, adjusted by 1 N

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HCl /NaOH). The effect of ion chelators (1-20 mM Ethylenediaminetetraacetic acid, EDTA;

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Sigma, U.S.A.) in reversing the effect of ion addition during gelation was studied in 4 mM CaCl2

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containing 3.0 % (w/v) SF blends (BM3.0:AA3.0, 1:1), wherein blends without Ca2+/EDTA and

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blends with only Ca2+ (4 mM) were used as positive (+ve) and negative (-ve) control, respectively.

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The effect of molecular weight of BM in gelation was studied by obtaining different molecular

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sized BM obtained by altering the degumming time, as described previously 45. Results obtained

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from turbidity, fluorescence and dynamic light scattering (DLS) analysis inferred that BM SF

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degummed at time points 5, 30 and 60 min were found to be as high molecular weight BM

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(hMWBM), medium molecular weight BM (mMWBM) and small molecular weight protein

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(sMWBM), respectively 45. 3 % (w/v) AA SF was blended with 3 % (w/v) BM SF of different

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molecular weights in the ratio of 1:1 and absorbance at 600 nm was recorded (37 °C). 3.0 % (w/v)

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each of hMWBM, mMWBM and sMWBM, and 3 % (w/v) AA SF (AA3.0) were used as control.

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To confirm the hydrophobic interaction between two types of protein, BM SF was replaced by 8 ACS Paragon Plus Environment

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bovine serum albumin (BSA; Sigma, U.S.A.) in the blends. 3 % (w/v) BSA was blended with 3 %

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(w/v) AA SF in the ratio of 1:1 and absorbance at 600 nm was recorded at 37 °C. BM1.5/AA1.5

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blended protein was used as positive control, whereas BSA3.0 (i.e., 3 % w/v BSA), AA3.0 and

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BM3.0 were used as negative controls.

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2.3.2. Intrinsic fluorescence spectra of sol-gel transition

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The conformational transformation in blended SF solution during the sol–gel transition was found

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out spectrofluorometrically (FluoroMax-4, Horiba) at 37 °C at excitation/emission wavelengths of

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280/600 nm. The intrinsic tryptophan (Trp) residues of SF were used as probe in fluorescence

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measurement. Scan speed was set at 1200 nm/min with an excitation/emission slit of 5 nm.

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Fluorescence emission spectra were collected at a resolution of 1 nm. The emission spectra were

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recorded at every 5 min up to 30 min for each sample. The maximum peak at 340 nm was

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normalized to compare gelation pattern among different blends. The full width half maxima

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(FWHM) at 340 nm of a blended SF was plotted to understand the change in fluorescence intensity

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per unit of time interval.

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2.4. Physicochemical characterization of hydrogels

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2.4.1. Field Emission Scanning Electron Microscopy (FESEM) and Energy-dispersive X-ray

17

spectroscopy (EDX)

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Surface morphology of the lyophilized hydrogels was studied using FESEM (Zeiss Sigma, Carl

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Zeiss AG, Germany) at an operating voltage of 2 kV. Hydrogels were frozen rapidly in liquid

20

nitrogen (-196 °C) and samples were freeze fractured for observation. The freeze fractured

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hydrogels were lyophilized followed by gold (Au) coating and images were captured under

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vacuum dried condition. Average pore size of each hydrogel was determined using ImageJ 1.4 9 ACS Paragon Plus Environment


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(Wayne Rasband) software by measuring ~50 pores, randomly. EDX analysis of hydrogels was

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conducted to measure the presence of any toxic elements e.g., Li/Br during silk fibroin isolation.

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2.4.2. Fourier Transform Infrared Spectroscopy (FTIR)

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The secondary structure of the hydrogels was analyzed using FTIR (Nicolet iS 10, Thermo

5

Scientific, U.S.A.). Briefly, the hydrogel samples were lyophilized followed by making of KBr

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disks in the ratio of 9:1 (KBr to sample). All absorbance spectra were recorded within the spectral

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range of 4000-400 cm-1 with a resolution of 4 cm-1 and 32 scans. Lyophilized powder of AA3.0

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and BM3.0 were used as control. Amide I region (1600 – 1700 cm-1) was examined to identify the

9

secondary structures in the hydrogels by performing a second order derivitization of the spectra

10

using OriginPro 8.5 software and finally the Lorentzian curve-fitting method was applied to the

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deconvoluted spectra. The absorbance bands in the range of 1650-1658 cm-1, 1620-1640 cm-1,

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1660-1680 cm-1 and 1670-1695 cm-1 were attributed as α-helix, β-sheets, loops, and anti-parallel

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β-sheets, respectively 46.

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2.4.3. Wide Angle X-Ray Diffraction (WAXD)

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The X-ray diffractogram of different sets of hydrogels were recorded with a Bruker D2 Phaser

16

diffractometer (U.S.A). Ni filtered CuKα was used as the X-ray source to investigate the changes

17

in crystallinity upon gelation. The diffraction range (2θ) was set between 10°-60° with a scan speed

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of 0.5° min-1. Powder forms of AA3.0 and BM3.0 were used as control.

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2.4.4. Swelling study

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Swelling behavior of hydrogels was carried out by immersing them into distilled water. Briefly,

21

dry weight of the lyophilized hydrogels was measured as (Wd) using an electronic balance


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(Sartorius, BSA224S-CW) followed by aqueous immersion at room temperature (25 °C) for 24 h.

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At definite time intervals, the swollen hydrogels soaked with water were carefully taken out and

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the wet weight (Ws) was recorded. The excess water surrounding the constructs was gently wicked

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out using tissue paper. The swelling ratio was expressed via the following equation 47.

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Swelling ratio = [(Ws – Wd)/Wd]

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2.4.5. In vitro enzymatic degradation

7

In vitro enzymatic degradation was performed to determine the stability of hydrogels in

8

physiological conditions. For this purpose, protease XIV from Streptomyces griseus (Sigma,

9

U.S.A.) with an activity of 3.5 U.mg-1 was used as described previously 47. In brief, different sets

10

of hydrogel were immersed in 3 ml of phosphate buffered saline (PBS, pH 7.4) containing 2 U.mg-1

11

of enzyme at 37 °C to achieve the optimum enzyme activity. The enzyme solution was replaced

12

with freshly prepared solution at every 3rd day. For control studies, hydrogels were also immersed

13

in PBS (pH 7.4) without enzyme under similar experimental conditions. Hydrogels were taken out

14

every 7th day, rinsed with distilled water, lyophilized, and dry weight was recorded to calculate the

15

loss over a period of 28 days.

16

2.4.6. Mechanical behavior of hydrogels

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Unconfined compressive strength of various sets of hydrogels in different physicochemical

18

conditions (e.g., various diameter of hydrogel, % of strain rate, time frame, pH and ionic strength)

19

was determined using Instron 5944 (Norwood, MA, U.S.A.), equipped with a 100 N load cell.

20

Hydrogels were cut into cylindrical shape (10 x 10 mm) and used for compressive test 1 h post

21

gelation. All tests were accomplished with an unconfined configuration and a displacement control

22

mode at a rate of 5 mm/min. After the compression tests, the stress and strain curves were plotted 11 ACS Paragon Plus Environment


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based on the measured cross-sectional area and sample height. Compressive strain was applied up

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to 80 % or until the construct got fractured. To mimic the native biomechanical condition of NP,

3

compressive modulus of all hydrogels of two different diameters (12 mm and 20 mm) with same

4

height (10 mm) was calculated for three different compressive strains i.e., 10 %, 20 % and 30 %.

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To determine the impact of storage time on mechanical properties, compressive modulus was also

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recorded after 1 h, 24 h and 14 day of gelation. All sets of hydrogels were incubated in PBS till

7

the aforementioned time points (pH 7.4) at 37 °C before testing. To mimic different physiological

8

conditions, hydrogels were immersed in three buffer systems having different pH (citrate buffer,

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pH 4.0; phosphate buffer, pH 7.4 and glycine-NaOH buffer, pH 9.0) and ionic strength (100, 500

10

and 1000 mM of NaCl) for 1 h prior to testing. To determine the recovery of mechanical properties,

11

hydrogels were transferred from both pH 4.0 and pH 9.0 to pH 7.0 following 1 h of incubation in

12

respective pH buffer systems.

13

A confined compressive modulus was also measured to mimic the confined NP tissue construct.

14

A manually prepared cylindrical plunger was compressed upon the hydrogel samples using a strain

15

rate of 5 mm/min. The hydrogel was confined by a high-density polypropylene ring. The

16

polypropylene ring was considered to be a harder body, as its modulus (approximately 10.0 GPa)

17

is several orders of magnitude higher than that of hydrogels. Confined compressive modulus was

18

recorded for each hydrogel at 10 % compressive strain. Other sets of hydrogel were used for

19

unconfined compressive testing as control.

20

For cyclic mechanical testing, hydrogels were compressed to axial strain of ~15 % for 50 cycles

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at a loading and unloading rate of 5 mm/minute in PBS solution (pH 7.4) at 37 °C. The axial strain

22

of 15% was selected as it is close to the actual strain value corresponding to loading on the


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complete intervertebral discs (IVD) in normal physiological conditions. The maximum

2

compressive stress was calculated for each complete cycle.

3

2.4.7. Rheological properties of hydrogel

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Dynamic oscillatory test for gelation point analysis, amplitude sweep for viscoelastic region (LVR)

5

and frequency sweep for frequency dependent viscoelastic properties of hydrogels were performed

6

on a Rheometer (MCR 302, Anton Paar) using a stainless steel parallel plate of 25 mm diameter

7

and a gap distance of 0.5 mm. The normal force on the sample was set to 0.1N during the entire

8

experimental procedure. To determine the gelation point, blended silk solution of different ratios

9

were placed in between two plates maintaining the temperature at 37 °C and data was collected at

10

a low strain amplitude (γ = 1 %, ω = 10 rad/s) to prevent possible sample manipulation due to

11

shear induced gelation in SF solution. A low viscosity mineral oil was applied surrounding the

12

plates to minimize the evaporation rate. Complex viscosity (η*) for each hydrogel system was

13

plotted against time. A strain amplitude sweep (1 % to 1000 %) was performed to determine the

14

LVR for each hydrogel system. The LVR of hydrogel was important for frequency sweep testing

15

in which elastic modulus (G′) and viscous modulus (G′′) were independent of strain amplitude.

16

Frequency sweep for each hydrogel system was collected over a wide frequency range (γ = 1 %,

17

ω = 1–100 rad/s). For thixotropic analysis, strain sweep was performed in the range of 1-1000%.

18

Step strain was set at 1000 % for 50 sec followed by 1% strain for 200 sec for three different sets

19

of hydrogel systems (BM0.5/AA2.5, BM1.5/AA1.5 and BM2.5/AA0.5).

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2.5. Biological studies on hydrogels

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2.5.1. Isolation and culture of porcine NP cells



1

Nucleus pulposus (NP) cells were isolated from porcine intervertebral disc (IVD) procured from

2

local abattoir within 1h post processing. In brief, IVDs from lumber portion were preferred and

3

adjacent muscle and tendons were discarded to separate the vertebral discs from the spine. The

4

entire dissection procedure was performed under sterile condition. The central gelatinous NP was

5

separated from the adjoining annulus fibrosus (AF) tissue. Thereafter, the isolated NP tissue was

6

thoroughly washed with PBS (pH 7.4) followed by digestion in 0.03 % collagenase (Sigma, St.

7

Louis, MO) in complete DMEM/F12 medium for 7-8 h. The digested slurry was then centrifuged

8

at 235 g for 5 min followed by 1-2 times PBS washing to remove extra enzyme fraction. The cell

9

number and viability were determined by trypan blue staining. NP cell pellets were resuspended

10

and plated at a density of 2 x 105 cells.cm-2 and placed at 37 °C in an incubator with 5 % CO2 and

11

85 % humidity. The culture medium [DMEM/F12 supplemented with 10 % Fetal bovine serum

12

(FBS), 100 U.ml-1penicillin G (Gibco, Life Technologies, U.S.A.), and 100 mg.ml-1 streptomycin

13

(Gibco, Life Technologies, U.S.A.)] was changed every 3rd day. Cells were passaged up to P4 and

14

stored at -196 °C for future use.

15

2.5.2. Cell seeding and proliferation assay

16

Cell proliferation was monitored using alamar blue dye reduction assay (Invitrogen, Life

17

Technologies, U.S.A.) at specified intervals of 1, 3 and 7 days. Equal number of NP cells (2 x 104

18

cells per hydrogel surface, 12 mm in diameter) was seeded on preconditioned hydrogels (complete

19

DMEM/F12 for 24 h) in a 24 well plate. Three different sets of hydrogels (BM0.5/AA2.5,

20

BM1.5/AA1.5 and BM2.5/AA0.5) were used for cell seeding purpose. Equal number of cells was

21

encapsulated in the hydrogels. The cell seeded hydrogels were incubated at 37 °C with 5 % CO2.

22

The medium was changed every alternative day. The cell proliferation was plotted in terms of

23

Alamar blue dye reduction. The absorbance was measured at 570 nm and 600 nm using a 14 ACS Paragon Plus Environment

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1

multiplate reader (Tecan Infinite 200 PRO series) and proliferation was calculated following the

2

manufacturer’s protocol. The cell encapsulated hydrogels were further used for cell viability

3

assessment.

4

2.5.3. Assessment of viability and arrangement of seeded cells in hydrogel

5

Viability of encapsulated NP cells in hydrogels was screened using live/dead assay kit (Invitrogen,

6

Life Technologies, U.S.A.) following the manufacturer’s protocol. In brief, live/dead assay reagent

7

was added to the cell loaded hydrogels and incubated at 37 °C for 15 min. Thereafter, hydrogels

8

were thoroughly washed with PBS (pH 7.4) and images were recorded using a florescence

9

microscopy (Evos FL, Life Technologies, U.S.A.).

10

2.5.4. Ex-vivo biomechanical study of hydrogel

11

For bio-functional assessment, blended SF solution was injected into the degenerated native disc,

12

allowed to form hydrogel and thereafter, cyclic mechanical testing was performed. For this

13

purpose, three different sets of native porcine lumber discs were collected from local abattoir

14

within 1h of post processing. Set I was used as control without any treatment. Set II and III were

15

used as degenerative discs. To degenerate, 1 ml mixture of collagenase (0.5 mg.ml-1) and protease

16

(2 U.ml-1) was injected into the central NP portion and incubated at 37 °C for 24 h in sterile

17

condition. After 24 h of degeneration, set II was injected with PBS and set III was injected with

18

blended SF solution (BM0.5/AA2.5) followed by 1 h incubation at 37 °C. Cyclic mechanical test

19

was performed with all three sets. All discs were compressed to the axial strain of ~2 % for 5

20

cycles at a loading and unloading rate of 1 mm/minute in PBS solution (pH 7.4) at 37 °C. The

21

stress-strain curve was plotted for all the cases.

22

2.6. Statistical analysis 15 ACS Paragon Plus Environment


1

All quantitative experiments were performed in triplicates and results are expressed as mean ±

2

standard deviation. Statistical analysis of data was performed by one way analysis of variance

3

(ANOVA) with Origin 6.0. Differences between groups of p 64%).

4

Crystallinity in hydrogel was further confirmed by WAXD. Peaks at 2θ=21° (major peak), 24°

5

(minor peak) and a shoulder peak at 2θ=42° confirmed the crystallinity in hydrogels (Figure 4C).

6

However, the controls BM3.0 and AA3.0 did not show β-sheet transition. EDX was utilized to

7

analyze the elemental or chemical compositions of the hydrogel samples under investigation.

8

Hydrogels did not contain any residual amount of LiBr ions, used during degumming of protein

9

fibers (Figure S8).



1 2 3

Figure 4. Physicochemical characteristics of hydrogels; (A) FESEM images of hydrogels: (I)

4

BM0.5/AA2.5, (II) BM1.0/AA2.0, (III) BM1.5/AA1.5, (IV) BM2.0/AA1.0, (V) BM2.5/AA0.5

5

and (VI) 200 KX magnification of hydrogel, (B) FTIR spectra, (C) WAXD analysis, (D) swelling 26 ACS Paragon Plus Environment

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1

study and (E) degradation study; mass remaining of hydrogel. Values are plotted as mean ±

2

standard deviation, n=3, where *** p ≤ 0.001, ** p ≤ 0.01 and * p≤ 0.05. For swelling study;

3

different letters (a, b, c, d and e) represent that the groups are significantly different from each

4

other (p ≤ 0.01).

5

3.9. Swelling study

6

Swelling ability of any construct applied in tissue engineering purpose is one of most important

7

necessities as it reveals the water retention capacity of that construct. To evaluate this intrinsic

8

property, lyophilized hydrogels were submerged in deionized water over a period of 24 h (Figure

9

4D). A sharp increase in swelling ratio was observed in all types of hydrogels within 20 min except

10

in BM2.0/AA0.5 blend which showed minimum swelling. A gradual yet slow increase was shown

11

up to 250 min and thereafter attained an equilibrium state. Swelling ratio increased with the

12

increase in the concentration of AA SF blends. Maximum swelling occurred in BM0.5/AA2.5

13

which was significantly higher than all the groups (p≤0.01), whereas BM2.5/AA0.5 showed

14

minimum swelling ratio. The swelling ratio of all the hydrogel groups was significantly different

15

(p≤0.01) from each other.

16

3.10. Degradation study

17

In tissue engineering, the applied biomaterials should get degraded with time and regeneration of

18

neo-tissue will seal the gap thus formed. To mimic the physiological biodegradable

19

microenvironment, hydrogels were treated with protease solution over 28 days (Figure 4E). The

20

degradation rate was represented as percent of mass remaining at each time point. There was a

21

significant difference in degradation rate of the hydrogels treated with enzymes group than that of

22

the control group (p≤0.001). It was observed that the degradation rate was decreased with the 27 ACS Paragon Plus Environment


1

increase of BM SF in the blends. BM2.5/AA0.5 experienced minimum degradation (~60 % mass

2

remaining), followed by BM2.0/AA1.0 (~38 % mass remaining) after a period of 28 days (p≤0.01).

3

However, there was no significant difference in degradation rate among BM1.5/AA1.5,

4

BM1.0/AA2.0 and BM0.5/AA2.5 blends, but significantly different from BM2.0/AA1.0 (p≤0.05)

5

and BM2.5/AA0.5 (p≤0.01).

6

3.11. Mechanical properties of hydrogel

7

Mechanical characterization of silk hydrogel is obligatory as it would be used to study the load

8

bearing applications like NP tissue engineering for disc degeneration therapy. The mechanical

9

properties of the hydrogels were analyzed under different in vitro conditions e.g., different time

10

frames post gelation, diameters, strain value, ionic strength, pH and mechanical confinement to

11

mimic the physiological microenvironment. The compressive modulus was calculated from stress-

12

strain curve for each hydrogel system in unconfined method at 10 % strain.

13

The compressive modulus of various hydrogels was calculated post 1 h of gelation. It was observed

14

that the compressive modulus was increased to a certain limit with the increase in concentration

15

of BM SF (Figure 5A). The maximum modulus value was obtained from BM2.0/AA1.0

16

(21.85±0.85 kPa), followed by BM1.5/AA1.5 (17.2±2.21 kPa), BM1.0/AA2.0 (11.37±1.07 kPa)

17

and BM0.5/AA2.5 (4.7±0.24 kPa). BM0.5/AA2.5 showed minimum compressive modulus which

18

was significantly lower than any other group studied (p≤0.01). However, BM2.5/AA0.5 did not

19

show any significant difference in compressive modulus from BM1.0/AA2.0, but showed

20

significantly less value than both BM1.5/AA1.5 and BM2.0/AA1.0 (p≤0.01).


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1 2 3

Figure 5. Mechanical properties of hydrogel; (A) compressive modulus post 1 h of gelation, (B)

4

diameter dependent compressive modulus measurement, (C) Compressive modulus at different

5

strain, and (D) time dependent compressive modulus measurement, (E) effect of pH on

6

compressive modulus, (F) effect of salt concentration on compressive modulus and (G) confined

7

mechanical testing of hydrogel: (I) schematic diagram of confined-unconfined mechanical testing,

8

and (II) comparative analysis of confined and unconfined compressive modulus of hydrogels.



1

Values are plotted as mean ± standard deviation, n=3, where # p≤ 0.001 vs Day 1, *** p ≤ 0.001,

2

** p ≤ 0.01 and * p ≤ 0.05.

3

As NP diameter varies from patient to patient, we used two different diameters of hydrogels; 12

4

mm and 20 mm. The compressive modulus of all hydrogel systems were significantly affected by

5

diameter. It was observed that as the diameter was increased, the compressive modulus was

6

increased as well (Figure 5B). For instance, 20 mm diameter of BM2.0/AA1.0 resulted in

7

compressive modulus of 48.6±3.9 kPa which was ~2.5 times higher than that of 10 mm diameter

8

of the same (p≤0.001).

9

The IVD can compress ~30 % of its vertical axis 48. To mimic such biomechanical conditions, the

10

compressive modulus of the hydrogels was calculated in three different strain rates, 10 %, 20 %

11

and 30 %. In all cases, the value got increased as the strain rate was increased (Figure 5C). For

12

instance, BM2.0/AA1.0 blend showed maximum compressive modulus (80.42±4.21 kPa) at 30 %

13

strain which was significantly higher than 20 % strain (49.61±5.32 kPa) (p≤0.001) and 10 % strain

14

(20.8±2.78) (p≤0.01), respectively.

15

The stability of hydrogels was assessed by maintaining them in PBS (pH 7.4) at three different

16

time frames; 1 h post gelation, 1 day post gelation and at day 14 (Figure 5D). The compressive

17

modulus of all types of hydrogels (except BM2.5/AA0.5) was not affected by a short incubation

18

time of 1 day, but significantly decreased after 14 days (p≤0.001). For instance, the obtained

19

compressive modulus of BM2.0/AA1.0 was 23.29±0.89 kPa on day 1, which was decreased to

20

13.83±0.26 kPa on day 14. In an exceptional case, BM2.5/AA0.5 showed maximum compressive

21

modulus of 12.42±1.21 kPa following 1 h of gelation, which was significantly increased to

22

23.22±1.35 on day 1 (p≤0.001) but decreased to 6.76±1.15 kPa on day 14.


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1

To understand the effect of physiological fluid environments on mechanical behaviors, hydrogels

2

were treated in three different pH (4.0, 7.4 and 9.0) and salt strengths (100, 500 and 1000 mM

3

NaCl). It was noticed that pH played a significant role on compressive modulus. In all the cases,

4

the compressive modulus was significantly increased after transferring the hydrogels from pH 7.0

5

to pH 4.0, whereas the value decreased when transferred to pH 9.0 (Figure 5E). For instance, the

6

compressive modulus of BM2.0/AA1.0 was significantly increased to 29.65±0.97 kPa from

7

23.29±0.89 kPa when transferred at pH 4.0 from pH 7.4 (p≤0.001), whereas the value decreased

8

to 17.77±0.67 kPa after transferring at pH 9.0 (p≤0.001). However, after transferring the hydrogels

9

from pH 7.4 to pH 4.0, compressive modulus increased by ~30 % in all cases except BM0.5/AA2.5

10

that showed ~80 % increase (Figure S9, A). Similarly, after transferring into pH 9.0, maximum %

11

decrease of compressive modulus was recorded for BM1.0/AA2.0 (~45 %), whereas the least value

12

was obtained for BM2.5/AA0.5. However, >30% decrease in compressive modulus was recorded

13

for the rest of the hydrogel systems (Figure S9, B). To check the recoverability of mechanical

14

properties, hydrogels were transferred from pH 9.0 to pH 7.4 and the compressive modulus was

15

measured. All the hydrogel systems had the ability to recover their compressive modulus by >70

16

% (Figure S9, C). However, change in salt concentration (100-1000 mM) did not show any

17

significant effect on compressive modulus (Figure 5F).

18

Gelatinous NP is confined by AF tissue. To mimic such microenvironment, compressive modulus

19

was conducted under confined system. The results showed that compressive modulus for all the

20

cases were significantly higher when measured under confined strategy than the unconfined way

21

(p ≤ 0.001) (Figure 5G, II). For instance, the unconfined compressive modulus of BM2.0/AA1.0

22

was 21.85±0.86 kPa, whereas the confined compressive modulus was 259.94±30.88 kPa (p≤0.001)

23

which was approximately 11 fold higher than the unconfined compressive modulus. 31 ACS Paragon Plus Environment


1

3.12. Cyclic mechanical study of hydrogels

2

The IVD is more often experienced to display continuous compression-decompression cycles

3

throughout life-time. Therefore, it is of utmost importance to perform the cyclic compression test

4

with NP analog of SF hydrogels. Cyclic compression test was conducted to study the regaining of

5

mechanical properties of all hydrogel systems (Figure S10A). Figure S10B, I shows the typical

6

stress-strain curve for all types of hydrogels. BM0.5/AA2.5 and BM1.0/AA2.0 did not show any

7

fissure even after applying 80 % strain. BM1.5/AA1.5 was broken down at 70 % strain, whereas

8

BM2.0/AA1.0 and BM2.5/AA0.5 showed permanent deformation at 40 % and 25 % strain,

9

respectively. From cyclic compression test, it was noticed that the area under stress-strain curve

10

gradually decreased after each cycle for all types of hydrogels (Figure S10B, II-V). It was

11

observed that maximum compressive stress at any particular strain decreased faster for initial

12

cycles and thereafter the rate of energy differences decreased slowly when 50 cycles were reached.

13

The maximum variation in compressive stress at 15 % strain from cycle 1-50 was recorded for

14

BM2.0/AA1.0, whereas BM0.5/AA2.5 showed minimum changes. BM0.5/AA2.5 was shown to

15

retain the maximum compressive stress of ~78 %, whereas this ability was gradually decreased as

16

the concentration of BM SF increased in the blends. For instance, BM2.0/AA1.0 showed minimum

17

recoverability of compressive stress (Figure S10B, VI)

18

3.13. Rheological properties of hydrogel

19

Rheological analysis was undertaken to study the flow properties or deformation of the hydrogels

20

at specific forces. Gelation point was determined in time sweep test of blended solutions. All the

21

blended solutions showed a typical sol feature (G”>G’) at initial time points. However, after

22

certain point, G’ crossed G” resulting in gel formation. It was observed that BM1.5/AA1.5 took at


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1

least (~270 s) time to form gel, whereas maximum time (~765 s) was taken by BM2.5/AA0.5. The

2

gelation time for BM0.5/AA2.5, BM1.0/AA2.0 and BM2.0/AA1.0 was ~610 s, ~300 s and ~500

3

s, respectively (Figure 6A). These results complimented the values obtained from absorbance or

4

fluorescence based assays for gelation study. The maximum storage modulus (~8686 Pa) was

5

gained by BM1.5/AA1.5 which was related to the faster crosslinking in the blend. Gelation is

6

directly related to the increase in complex viscosity (η*) with time. Here, η* was gradually

7

increased with time for all cases supporting the gelation occurrence (Figure 6B).

8 9

Figure 6. Rheology of SF hydrogel; (A) Time sweep for gelation point analysis, (B) complex

10

viscosity vs. time, (C) Determination of LVR for hydrogels, (D) frequency sweep and (E)

11

thixotropic test

12

In amplitude sweep, shear strain (γ) is varied while frequency is kept constant. For analysis, G’

13

and G” were plotted against γ. At low γ, G’ and G” were constant for all cases. As the γ value

14

increased, deformation was started leading to decrease in G’ value that crossed G”. The slope value

15

of G’, known as linear viscoelastic region (LVR) was gradually decreased as the concentration of 33 ACS Paragon Plus Environment


1

BM SF increased in the hydrogels (Figure 6C). BM 0.5/AA2.5 displayed the maximum value and

2

LVR extended up to ~222 % strain, whereas BM2.5/AA0.5 showed the least strain value (γ = ~14

3

%) for LVR. A moderate value (γ = ~165 %) was obtained for BM1.5/AA11.5. In frequency

4

sweep, frequency is varied while γ remains fixed. All the hydrogels showed no deformation up to

5

100 rad/s. G’ and G” were frequency independent for all hydrogels (Figure 6D). In the thixotropic

6

analysis, it was observed that all the hydrogels had the recoverability property from high strain

7

induced deformations. However, the greater recoverability was shown by BM 0.5/AA2.5 (Figure

8

6E).

9

3.14. Biological assessment

10

Cytocompatibility assessment of any biomaterial used for tissue engineering applications is one of

11

the most important pre-requisites. For this purpose, primary NP cells, isolated from porcine source,

12

were seeded on hydrogel surface and cultured for 7 days (Figure 7A). Alamar blue reduction assay

13

was performed to assess the proliferation rate of the primary cells upon contact with hydrogel

14

surface. It was observed that NP cells proliferated over 7 days on all types of hydrogels surface.

15

However, maximum number of cell proliferation occurred on BM0.5/AA2.5 which was

16

significantly higher than that of BM1.5/AA1.5 or BM2.5/AA0.5 on day 7 (p ≤ 0.01) (Figure 7B).

17

To check the cellular viability inside hydrogel microenvironment, NP cells were encapsulated

18

within hydrogels and maintained for 7 days in DMEM/F12 (Figure 7C, I-III). Live cells were

19

visualized using live/dead assay kit. Cells were viable in all hydrogel systems, but enhanced green

20

fluorescence signal was recorded for BM0.5/AA2.5.


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1 2 3

Figure 7. Biological assessment of silk hydrogel; (A) isolation of NP hydrogel and NP cell form

4

porcine vertebral disc: (I) native vertebral disc, (II) native NP gel and (III) cultured NP cells on

5

tissue culture plate, (B) Cell proliferation assay using Alamar blue reduction method and, (C)

6

viable cells (green fluorescence) inside the hydrogels: (I) BM0.5/AA2.5, (II) BM1.5/AA1.5 and

7

(III) BM2.5/AA0.5. Data are plotted as a mean ± standard deviation, n=3, where ** p ≤ 0.01.

8

3.15. Ex-vivo biomechanical study of hydrogel

9

Ex-vivo biomechanical study was performed by injecting the blended SF solution into the

10

degenerated NP portion of disc. Injected solution was allowed to form hydrogel inside the disc.

11

Three sets were used for this study; normal disc (control group), degenerated disc filled with PBS



1

and hydrogel assisted regenerated disc (Figure 8, A-B). Cyclic compression test was performed

2

with all the three types of discs. In control group (Set I), stress-strain curve area remained

3

unchanged throughout the cyclic process (Figure 8C, I). The PBS filled degenerated disc (Set II)

4

showed drastic drop in maximum compressive stress after 5 cycles (Figure 8C, II). However, the

5

hydrogel assisted regenerated disc (Set III) maintained its compressive properties to a satisfactory

6

level (Figure 8C, III).

7


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1

Figure 8. Ex-vivo biomechanical study; (A) three sets of porcine vertebral discs, (B) injection of

2

blended SF solution into the disc, (C) cyclic compressive test: (I) native disc, (II) degenerated disc

3

and (III) hydrogel assisted regenerated disc, (D) three sets of dissected discs after test.

4

4. Discussion

5

Silk hydrogels have been used over the years in different tissue engineering applications including

6

cartilage, bone, and intervertebral disc

7

into hydrogel by inducing structural transition into β-sheets utilizing several physiochemical

8

methods

9

expenditure of hydrogel fabrication, but also increase the chances of chemical toxicity. However,

10

the mechanical properties of fabricated hydrogels are also unsatisfactory, specifically when

11

applied in load bearing applications like NP tissue engineering. Therefore, there is an urgency need

12

for a facile and rapid gelling technique with tunable mechanical and bioresponsive properties,

13

avoiding the usage of any hazardous chemicals.

14

NP is the gelatinous central tissue of vertebral disc. It is confined by multiple sheets of AF tissue

15

and dissipates the pressure to this surrounding AF. Once degeneration of the disc starts it fails to

16

regenerate as it is an avascular organ. A minimally invasive way is to inject the hydrogel into the

17

damaged disc via in situ gelation approach that may overcome the surgically associated treatment

18

fallacy. In the present study, we blended two different SF (B. mori and A. assamensis) to make

19

hydrogel independent of any external stimuli. SF from two different sources self-assembled with

20

time forming hydrogel. This self-assembly might be due to the alteration in hydrophobic-

21

hydrophilic microenvironment. The differential hydrophobicity that caused the proteins to interact

22

happened because of the amino acid composition, sequence and arrangement of the protein chain.

24.

40, 49.

The aqueous SF solution can easily be transformed

These time consuming, labor-intensive methods do not only increase the final



1

The primary structure of BM SF mainly consists of AGSGAG repeats which comprises ~55% of

2

the total fibroin. These repeats are mainly involved in crystalline structure formation which in turn

3

are primarily responsible for conformational changes of protein (silk I to silk II) when subjected

4

to bio-spinning or under laboratory conditions (e.g., alcohol treatment, sonication, pH and ionic

5

strength variations)

6

and rest are 5.5% tyrosine, 1% aspartic and glutamic acid, and 2.2% isoleucine, leucine and valine

7

43.

8

serine and 5.5% tyrosine 43. There are mainly three motifs in AA SF secondary confirmations; A-

9

motif that has only alanine residue, G-motif mainly consisting of glycine residues and R-motif

10

containing arginine residues. Among these three motifs, A-motif is mainly responsible for

11

crystallinity as it contains poly-alanine stretches of (alanine)5-15.

12

The gelation was accelerated or decelerated when reaction occurred at high (42 °C) or low

13

temperature (25 °C), respectively (Figure 1A and 2A). This could be related to the hydrophobic

14

interaction between SF chains. It is a well-known phenomenon that hydrophobic interaction is

15

highly influenced by temperature and the strength of hydrophobic interaction increases as the

16

temperature raises 53. Here, two SF proteins of different hydropathicity (due to their amino acid

17

composition, sequence and arrangement) when blended, the protein chains were intended to

18

interact to each other due to the hydrophobic interaction. At the initial stage of gelation, the

19

hydrophobic regions in the protein chains began to rearrange themselves (inter-chain interaction)

20

which was further accelerated by intra-chain interaction as the gelation continued. The ratio of two

21

different hydrophobic chains also played a fundamental role in gelation. It was observed that when

22

two proteins were blended at a ratio of 1:1 (BM1.5/AA1.5), faster gelation occurred than any other

23

combinations at any given temperature. BM SF did not show any sign of gelation at any

45, 50-52.

The BMSF consists of 30.2% alanine, 45.9% glycine, 12.1% serine

Similarly, the primary structure of AA SF consists of 42.5% alanine, 28.9 % glycine, 10.2%


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1

temperature, whereas gelation initiation occurred in AA SF in the later stage either at 37 °C or at

2

a higher temperature. As AA SF itself possessed longer hydrophobic chains of poly-alanine,

3

change in temperature lead to the formation of self-assembled chains (Figure S1). Furthermore,

4

we conducted an experimental set up where a type of SF concentration got fixed while varying the

5

other type. It was observed that the protein combinations containing higher amount of AA SF,

6

showed faster onset of gelation (Figure 2B, II-III and Figure S2). This might be due to the

7

presence of longer hydrophobic fragments in AA SF chain that had more potential to self-

8

assemble, and BM SF is hypothesized to play the role of an inducer in gelation (Figure S1).

9

Gelation process was also associated with total protein concentration in the blends. It was noticed

10

that gelation time gradually decreased with the increase in total SF concentration in the blends

11

(Figure 2B, I). Higher amount of protein corresponds to higher probability of resulting in

12

development of hydrophobic chains which is correlated to the faster interaction that resulted in

13

faster gelation.

14

To understand the effect of various ions present in physiological fluids, gelation process was

15

conducted in presence of Ca2+ and K+ ions (1-4 mM). It was seen that gelation time gradually

16

increased with the increase in Ca2+ ions in the blend (Figure 2C, I). Previously, it was reported

17

that faster gelation occurred in presence of Ca2+ ions (20-200 mM) in sonication induced BM SF

18

solution 22. The high concentration of Ca2+ ions coagulated SF that finally led to the formation of

19

gel as a result of “salting out” effect. In the present study, we used a low range of Ca2+

20

concentration (1-4 mM) similar to the biological fluids. This significantly low ionic strength could

21

not induce coagulation or precipitation in BM SF rather induces “salting in” effect. In contrast,

22

AA SF was very sensitive to “salting out” effect even at a very less ionic strength (Video S1 and

23

Video S2). However, visible BM SF protein coagulation was observed at 6 mM Ca2+, and therefore 39 ACS Paragon Plus Environment


Page 40 of 68

1

it could be hypothesized that coagulation might start below the ionic strength level used in this

2

study (Figure S3, A). The hydrophobic interaction is directly correlated to salt concentration 54.

3

With the increase in salt concentration, it takes out the water molecule surrounding the protein

4

molecules keeping the hydrophobic residues close together. In this study, the gelation time was

5

increased in presence of Ca2+ ions. This might be due to the absence of hydrophobic chains in AA

6

SF, required to entangle with BM SF chains during gelation, as it was already precipitated or

7

separated from the blended solution. One possible reason behind the AA SF precipitation was the

8

presence of high content of Lys, Arg and His (3.9 %) that possess greater relative calcium-binding

9

strength over other amino acids (Figure S3, B)

55.

Study conducted by a group suggested that

10

higher Ca2+ ion concentrations maintain the random coil in BM SF

11

possible reason restricting one of the SF to play the role in hydrogel formation. However,

12

monovalent K+ ion did not show any effect on gelation process (Figure 2C, II). There were few

13

contradictory reports on the effect of K+ ions on SF gelation 22, 57. Therefore, it can be inferred that

14

the presence of Ca2+ in the physiological fluid hinders the occurrence of in situ SF gelation. For a

15

successful in situ gelation, EDTA was added to the blended SF solution and recoverability of

16

gelation time was checked (Figure 3A). It was observed that EDTA chelated the Ca2+ ions thus

17

setting the hydrophobic chains free for interaction.

18

The gelation time was also influenced by the molecular size of BM SF. The higher degumming

19

time provided low molecular weight BM SF and vice versa. Molecular size by virtue of the

20

hydrodynamic size was confirmed by conducting fluorescence and DLS analysis. It was noticed

21

that sMWBM (~15 nm in size) showed reduced fluorescence intensity at 340 nm than mMWBM

22

or hMWBM (Figure S4). The less degraded samples possessed higher tryptophan to tyrosine

23

fluorescence than the highly degraded samples. This was because of less probability of 40 ACS Paragon Plus Environment

56.

This might be another

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1

fluorescence resonance energy transfer (FRET) between the two residues in smaller peptide chains

2

of highly degraded sample 45. However, it was observed that faster gelation occurred in the blends

3

containing sMWBM (Figure 3B). This might be due to the ability of smaller BM SF chains to

4

penetrate into the core and interact faster with the relatively larger entangled AA SF chains (Figure

5

S5). In order to confirm the hydrophobic interaction between two SF types during gelation, we

6

replaced BM SF with BSA that does not have any hydrophobic motif (Figure 3C). It was observed

7

that blends with BSA showed the same gelation pattern as control AA3.0, indicating a time

8

dependent self-assembly phenomenon.

9

The intrinsic fluorescence of SF was also utilized to understand the gelation process. Fluorescence

10

intensity gradually increased as gelation process continued (Figure 3D). The protein molecules

11

were in Brownian motion when they were in liquid state. The gelation initiated when two protein

12

molecules started to interact with each other through hydrophobic interaction. The enhanced

13

fluorescence intensity with the gelation progression might be due to better fluorescence resonance

14

energy transfer between tryptophan to tyrosine residues when it was in close association within

15

hydrogel (cross-linked) microenvironment (Figure S6) 45. The maximum and faster increment of

16

fluorescence intensity was recorded for BM1.5/AA1.5 indicating faster interaction among the

17

chains when two proteins were in equal ratio (Figure 3D, II). To understand whether gelation was

18

an exponential/homogeneous or inhomogeneous process, we calculated the change in FWHM from

19

fluorescence intensity graph for each unit time interval and plotted them against the gelation time.

20

The zigzag pattern of change in FWHM depicted that the gelation progression was an

21

inhomogeneous or discontinuous process. We hypothesized if gelation was a homogeneous

22

process i.e., same number of both molecules (n+ n, n+ n, …….n+ n) interacted with each other at

23

each unit time, a straight line of change in FWHM would be obtained (blue dotted line). Similarly, 41 ACS Paragon Plus Environment


1

if gelation was an exponential process i.e., an increasing order of both molecules (n+ n, 2n+ 2n,

2

…….Nn+ Nn) interacted to each other; a linear pattern of change in FWHM against time would be

3

obtained (green dotted line). (Figure 3D, III).

4

From the FESEM data, it was noticed that the pore size of hydrogels gradually decreased as the

5

percentage of BM SF increased in the blends. Maximum pore size was obtained from

6

BM0.5/AA2.5, whereas BM2.5/AA0.5 showed minimum pore size (Figure 4A). Here, the smaller

7

hydrophobic fragments (GSGAGG) of BM SF chains (MW ~390 kDa) behaved as filler to this

8

hybrid polymeric network system. On the other hand, AA SF possess relatively longer chains

9

(~460 kDa) with greater hydrophobicity (poly-Ala) that interacted with themselves, leaving the

10

free spaces between the two zones of the interface. This event finally resulted in the formation of

11

bigger pores (Figure S8, A). This type of phenomenon was also observed in freeze dried scaffolds

12

from Antheraea mylitta SF

13

blends, it occupied the free spaces, thereby resulting in reduced pore size. The β-sheets transition

14

in the hydrogels was confirmed by FTIR (Figure 4B). The maximum amount of β-sheets was

15

observed in BM1.5/AA1.5. This might be due to the faster interaction between chains that finally

16

lead to the faster conformational changes. Gelation time can be correlated with the β-sheets content

17

of the hydrogels. It was evident from the β-sheets content (~83 %) of BM0.5/AA2.5 which showed

18

faster gelation than BM2.5/AA0.5 having the β-sheets content of ~65 %) (Table S6).

19

Swelling ability of hydrogels or scaffolds is one of the most important prerequisites for tissue

20

engineering applications. The swelling ratio increased with the increase of BM SF in the blend

21

(Figure 4D). This could be possibly because of the gross hydrophobicity and cross linking density

22

in the blends. BM0.5/AA2.5 showed minimum swelling owing to the maximum amount of

47.

However, as the concentration of filler BM SF increased in the


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1

hydrophobic chains it possess as well as high amount of β-sheets content (~83 %) that further

2

resisted the hydrogel from swelling.

3

Biodegradation of implanted materials should be accompanied with the neo-tissue regeneration.

4

For biodegradation purpose, a protease cocktail was used. It showed that blends with higher

5

amount of AA SF were degraded relatively at a faster rate (Figure 4E). This might be due to

6

presence of higher amount of target sites (Xaa-hydrophobic-cut-hydrophobic-Xaa, where Xaa is

7

any amino acid and hydrophobic amino acid is Ala) for proteases in AA SF chains.

8

Mechanical behaviors of any implantable system have to be analyzed to understand its efficacy

9

upon implantation, particularly in load bearing applications. The compressive modulus was

10

gradually increased as the concentration of filler BM SF increased in blends. The minimum

11

compressive modulus was recorded for BM0.5/AA2.5 in which longer hydrophobic chains of AA

12

SF were predominant (Figure 5A). Basically it was a homogenous complex of same molecules

13

that rearranged themselves in a systematic way. However, as the filler BM SF concentration

14

increased, the composite became a heterogeneous polymeric network. However, reduced

15

compressive modulus was recorded for BM2.5/AA0.5 as it became a homogeneous complex. It is

16

a well-known fact that mechanical strength is higher in heterogeneous matrix than homogenous

17

complex 58-59. If we compare between two “close to homogenous” complexes i.e., BM2.5/AA0.5

18

and BM0.5/AA2.5, the later combination showed less compressive modulus. This could be

19

correlated to the larger pores in BM0.5/AA2.5 leaving more void spaces making it more sponge

20

like material, whereas BM2.5/AA0.5 showed solid like construct comprising of higher mechanical

21

strength. From the stress-strain curve, it was evident that the maximum compressive stress was

22

gradually decreased with the increase of BM SF in the blends. Hydrogels with high amount BM

23

SF became more brittle than the hydrogels containing high amount AA SF. To understand the 43 ACS Paragon Plus Environment


1

deformation behavior of hydrogels, the cyclic compressive testing was performed. The hydrogels

2

with higher amount of AA SF (e.g., BM0.5/AA2.5) showed minimum loss of ultimate compressive

3

stress and retained ~80 % elasticity after 50 cycles (Figure S10B, VI). The high elasticity of

4

BM0.5/AA2.5 might be due to the longer chains of AA SF that entangled making a sponge like

5

matrix, whereas the filler BM SF filled the pores and became a solid like construct.

6

To understand the impact of geometry on mechanical properties, two different diameters of

7

hydrogels were selected; 12 mm and 20 mm. It was noticed that compressive modulus was

8

increased with the increase in diameter for any hydrogel systems (Figure 5B). This might be due

9

to the presence of more cross-linking (chemical or physical) in larger diameter group. There are

10

various reports that indicate the impact of geometrical parameter on compressive modulus of

11

hydrogel implants 48, 60. Therefore, the volume of implanted hydrogel might play an important role

12

in disc biomechanics. It could be expected that larger diameter of hydrogel might provide higher

13

compressive modulus allowing the intervertebral discs to withstand more loads. Moreover,

14

hydrogels in confined mechanical testing showed higher values than unconfined method for all

15

hydrogel systems. There are several factors (viz., under-filling, over-filling or line-to-line fit to the

16

nucleus cavity with hydrogel or confinement of NP) influencing the biomechanics that could be

17

resolved by this in vitro approach. To check the stability of the fabricated hydrogels, they were

18

placed in PBS (pH 7.4) and compressive modulus was recoded on day 1 and day 14. The

19

compressive modulus did not change significantly after day 1 of gelation for all cases, except in

20

BM2.5/AA0.5 where the compressive modulus was significantly increased (~2 fold). It could be

21

hypothesized that the gelation process in this combination was very slow and took the maximum

22

time to form hydrogel in which relatively smaller BM SF chains interacted with each other making

23

a highly compact construct (Figure 5D). 44 ACS Paragon Plus Environment

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1

To understand the effect of biological fluids on compressive modulus, hydrogels were treated in

2

solutions of different pH and ionic strength. The compressive modulus of all the hydrogels was

3

increased in pH 4.0, whereas the value was decreased in pH 9.0 (Figure 5E). The isoelectric pH

4

of silk fibroin is varied from pH 4.0-5.0 61. After transferring the hydrogels into the solution of pH

5

below the isoelectric point, intermolecular interaction increased inside the hydrogel resulting in

6

higher compressive modulus. In contrast, the higher pH (9.0) might degenerate the protein chains

7

that caused reduced compressive modulus. BM0.5/AA2.5 showed the maximum value of

8

percentage increase of compressive modulus in response to pH 4.0. It was therefore confirmed that

9

AA SF had greater pH sensitivity than BM SF. All the hydrogel systems recovered ~70 % of

10

compressive modulus after getting transferred from alkali pH (9.0) to neutral pH (7.4) (Figure

11

S9). However, the compressive modulus remained unchanged in salt concentration up to 1000 mM

12

NaCl (Figure 5F). From amplitude sweep of hydrogels, liner viscoelastic region (LVR) was

13

determined. LVR was extended till ~222 % of strain for BM0.5/AA2.5. This might be due to the

14

interpenetrating network of longer and homogeneous AA SF chains (Figure 6C) in the

15

BM0.5/AA2.5. Thixotropic test confirmed the reversibility of the hydrogel structures under very

16

high strain (1000 %).

17

Cytocompatibility assessment of a biomaterial used for tissue engineering applications is very

18

much essential for its clinical translation. All the hydrogels were found to be compatible for the

19

growth and proliferation of NP cells, enhanced proliferation being observed for the hydrogels

20

having higher amount of AA SF (Figure 7). This might be due to presence of RGD (arginine-

21

glycine-aspartic acid) motif in AA SF that supported faster cell-matrix adhesion and proliferation.

22

Few reports in the recent past have suggested this plausible effect of AA SF in various tissue

23

engineering applications 13-14, 62-63. Though the NP tissue exhibits fewer number of cells, it could 45 ACS Paragon Plus Environment


Page 46 of 68

1

be expected that the high proliferation rate would provide the sufficient amount of healthy cells

2

within a very short period of time and the healthy cells would secret adequate amount of ECM

3

molecules to recover the discs from disease condition.

4

Ex-vivo IVD organ culture model is a well-established technique to understand the biological and

5

biomechanical measures of disc health and disease

6

injectable hydrogel to be used as NP substitute to restore the disc biomechanics in ex-vivo model.

7

For instance, Showalter et al., used the degenerated human lumber discs wherein hydrogel (20 %

8

teleostean, 3 % N-carboxyethyl chitosan, and 7.5 % oxidized dextran) was injected to restore the

9

biomechanics after physiologic cyclic compressive loading in an ex-vivo model

64-66.

However, there are several reports on

67.

In a similar

10

kind of study, Malhotra et al., investigated the hyaluronic acid-gelatin based injectable hydrogel

11

to restore the range of motion of degenerated discs upon hydrogel implantation in the ex-vivo

12

model 68. In the current study, an ex-vivo model system was developed in order to speculate the

13

feasibility of these hydrogels in NP tissue engineering applications. The hydrogel assisted

14

regenerated disc showed good recovery of compressive modulus (Figure 8) when compared to

15

degenerated disc. So, the previous reports and our current study may provide a solid foundation to

16

evaluate the efficacy and reliability of ex-vivo model for better understanding of disc biomechanics

17

using hydrogel implants.

18

For in situ applications, the faster gelation always gets precedence. BM1.5/AA1.5 combination

19

took lesser time than any other combinations. On the other hand, compressive modulus of

20

BM05/AA2.5 was in the range of native NP gel (5.39 ± 2.56 kPa) 35. Other plausible factors of

21

BM05/AA2.5 included superior elasticity (Video S3) and cytocompatibility than rest of the

22

combinations. However, BM05/AA2.5 was highly sensitive to pH changes. Therefore, it could be


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1

a novel approach to select a suitable combination of silk hydrogels in a particular physiological

2

condition.

3

5. Conclusion

4

We have successfully developed the formulation of silk based in situ gelation technology towards

5

NP tissue engineering. Two naturally derived silk proteins of different hydrophobicity were self-

6

assembled independent of external stimuli and form hydrogel. The gelation process could be tuned

7

by types, concentration and molecular weight of SF, temperature, pH and ionic concentrations.

8

The tunable mechanical properties e.g., compressive modulus and elastic behavior were achieved

9

by varying the silk fibroin ratios. The efficiency of the hydrogels lies in the fact that they could

10

recover the compressive modulus from adverse conditions. Finally, the ex-vivo study supported

11

the possibility of hydrogels to its clinical translation.

12

Conflicts of interest

13

There are no conflicts to declare

14

Acknowledgements

15

B.B.M. greatly acknowledges the generous funding support from Department of Biotechnology

16

(DBT) and Department of Science and Technology (DST), Government of India. B.K.B.

17

acknowledges Ministry of Human Resource Development (MHRD), India for providing

18

scholarship. Dr. Aparajita Ghosh and Joseph Christakiran M are acknowledged for their generous

19

help and valuable suggestion in manuscript preparation. We also acknowledge Central Instrument

20

Facility (CIF) and Department of Biosciences and Bioengineering (BSBE), IITG for providing

21

high-end instruments.



1

Supporting Information:

2

Detailed information of gelation kinetics of silk blends in different conditions, different secondary

3

conformations of hydrogels constructs, schematic representation of hydrophobic interaction

4

among SF chains, effect of ionic concentration on gelation, proposed mechanism of molecular

5

weight (BM SF) effect on gelation, FRET during gelation and mechanism behind pore formation,

6

comparative gelation kinetic study between groups, effect of degumming time on molecular size,

7

deconvolution of FTIR spectra, mechanical properties of hydrogels in response to pH changes,

8

cyclic mechanical testing of hydrogels.

9 10 11 12 13 14 15 16 17 18 19 48 ACS Paragon Plus Environment

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1

For Table of Contents Use Only

2

Exploring Gelation and Physico-chemical Behavior of In-situ Bioresponsive Silk Hydrogels

3

for Disc Degeneration Therapy

4

Bibhas K. Bhunia and Biman B. Mandal*

5

Biomaterial and Tissue Engineering Laboratory, Department of Biosciences and Bioengineering,

6

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

7 8

Author for correspondence: *[email protected], *[email protected]

9



Figure 1. Silk hydrogel; (A) stepwise process of hydrogel preparation (B) schematic representation of gelation mechanism and (C) hydrogels of five different ratios (tube inversion method)

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Figure 2. Gelation pattern of silk hydrogel in different physiological conditions; (A) effect of temperature on gelation profile: (I) 25 °C, (II) 37 °C and (III) 42 °C, (B) effect of SF concentration and SF types on gelation: (I) concentration dependent gelation where total SF protein concentration varies from 1-4 %, (II) Group -1; blends of fixed AA concentration (2 %) with varying BM concentration (0.5-3.0 %) and (III) Group-2; blends of fixed BM concentration (2 %) with varying AA concentration (0.5-3.0 %), (C) effect of ionic concentrations on gelation: (I) effect of Ca2+ ions (II) effect of K+ ions and (III) effect of pH. Data represents mean ± SD (n = 3).



Figure 3. Gelation pattern of silk hydrogel; (A) reversal of Ca2+ ions effect on gelation using EDTA, (B) effect of different molecular weight of BM SF on gelation, (C) confirmatory assessment of hydrophobic interaction behind gelation and (D) intrinsic fluorescence analysis of silk hydrogel: (I) time dependent increase in fluorescence intensity of hydrogel, (II) comparative analysis of fluorescence intensity values among the blends and (III) Graph representing the changes in fluorescence intensities per unit of time: the blue and green dotted lines are the hypothetical situation of a homogeneous and exponential gelation process. Data represents mean ± SD (n = 3).


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Figure 4. Physicochemical characteristics of hydrogels; (A) FESEM images of hydrogels: (I) BM0.5/AA2.5, (II) BM1.0/AA2.0, (III) BM1.5/AA1.5, (IV) BM2.0/AA1.0, (V) BM2.5/AA0.5 and (VI) 200 KX magnification of hydrogel, (B) FTIR spectra, (C) WAXD analysis, (D) swelling study and (E) degradation study; mass remaining of hydrogel. Values are plotted as mean ± standard deviation, n=3, where*** p ≤ 0.001, ** p ≤ 0.01 and * p≤ 0.05. For swelling study; different letters (a, b, c, d and e) represent that the groups are significantly different from each other (p ≤ 0.01).



Figure 5. Mechanical properties of hydrogel; (A) compressive modulus post 1 h of gelation, (B) diameter dependent compressive modulus measurement, (C) Compressive modulus at different strain, and (D) time dependent compressive modulus measurement, (E) effect of pH on compressive modulus, (F) effect of salt concentration on compressive modulus and (G) confined mechanical testing of hydrogel: (I) schematic diagram of confined-unconfined mechanical testing, and (II) comparative analysis of confined and unconfined compressive modulus of hydrogels. Values are plotted as mean ± standard deviation, n=3, where # p≤ 0.001 vs Day 1, *** p ≤ 0.001, ** p ≤ 0.01 and * p ≤ 0.05.


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Figure 6. Rheology of SF hydrogel; (A) Time sweep for gelation point analysis, (B) complex viscosity vs. time, (C) Determination of LVR for hydrogels, (D) frequency sweep and (E) thixotropic test



Figure 7. Biological assessment of silk hydrogel; (A) isolation of NP hydrogel and NP cell form porcine vertebral disc: (I) native vertebral disc, (II) native NP gel and (III) cultured NP cells on tissue culture plate, (B) Cell proliferation assay using Alamar blue reduction method and, (C) viable cells (green fluorescence) inside the hydrogels: (I) BM0.5/AA2.5, (II) BM1.5/AA1.5 and (III) BM2.5/AA0.5. Data are plotted as a mean ± standard deviation, n=3, where ** p ≤ 0.01.


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Figure 8. Ex-vivo biomechanical study; (A) three sets of porcine vertebral discs, (B) injection of blended SF solution into the disc, (C) cyclic compressive test: (I) native disc, (II) degenerated disc and (III) hydrogel assisted regenerated disc, (D) three sets of dissected discs after test.



Table of Content (ToC) Graphic 367x281mm (96 x 96 DPI)


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Exploring Gelation and Physico-chemical Behavior of In-situ

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