Multiscale porosity in compressible cryogenically 3D printed gel for

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Applications of Polymer, Composite, and Coating Materials

Multiscale porosity in compressible cryogenically 3D printed gel for bone tissue engineering Deepak Gupta, Atul Kumar Singh, Ashwin Dravid, and Jayesh R. Bellare ACS Appl. Mater. Interfaces, Just Accepted Manuscript • DOI: 10.1021/acsami.9b05460 • Publication Date (Web): 13 May 2019 Downloaded from http://pubs.acs.org on May 13, 2019

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ACS Applied Materials & Interfaces

Multiscale porosity in compressible cryogenically 3D printed gel for bone tissue engineering Deepak Gupta1, Atul Kumar Singh2,3, Ashwin Dravid1,4, Jayesh Bellare*,1,2,5,6 1Chemical

Engineering Department, Indian Institute of Technology Bombay, Powai, Mumbai400076, India, 2Centre for Research in Nanotechnology & Science, Indian Institute of Technology Bombay, Powai, Mumbai-400076, India 3Central Research Facility (CRF), Indian Institute of Technology Delhi, New Delhi-110016, India, 4Johns Hopkins University, Chemical and Biomolecular Engineering 323 E 33rd St Baltimore, MD, 21218, USA, 5Tata Centre for Technology and Design, Indian Institute of Technology Bombay, Mumbai-400076, India, 6Wadhwani Research Centre for Bioengineering (WRCB), Indian Institute of Technology Bombay, Mumbai-400076, India. *Corresponding Author: Jayesh R. Bellare, Department of Chemical Engineering, Indian Institute of Technology Bombay, Mumbai-400076, India. Email: [email protected], Phone: +91 22 2576 7207, Fax: +91 22 2572 6895 or +91 22 2572 3480.

Abstract 3D printing technology has seen several refinements when introduced in the field of medical devices and regenerative medicines. However, it is still a challenge to 3D print gels for building complex constructs as per the desired shape and size. Here, we present a novel method to 3D print gelatin/carboxymethylchitin/hydroxyapatite composite gel constructs of complex shape. The objective of this study is to fabricate a bioactive gel scaffold with a controlled hierarchical structure. The hierarchy ranges from 3D outer shape to macro-porosity to micro-porosity and rough surface. The fabrication process developed here uses 3D printing in a local cryogenic atmosphere followed by lyophilization and crosslinking. The gel instantly freezes after extrusion on the cold plate. The cooling action is not limited to the build-plate but the entire gel scaffold is cooled during the 3D printing process. This enables the construction of a stable self-sustaining large-sized 3D complex geometry. Further, lyophilization introduces bulk micro-porosity into the scaffolds. The outer shape and macroporosity were controlled with the 3D printer, while the microporous structure and desirable rough surface morphology were obtained through lyophilization. With cryogenic 3D printing, up to 90% microporosity could be incorporated into the scaffolds. The microporosity and pore size distribution were controlled by changing the crosslinker and total polymer concentration, which gave as much as six times increase in surface

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open pores of size < 20 µm on increasing the crosslinker concentration from 25 mg/mL to 100 mg/mL. The introduction of bulk microporosity was shown to increase swelling by 1.8 times along with a significant increase in human umbilical cord mesenchymal stem cells (hUCMSCs) and Saos-2 cell attachment (2x), proliferation (2.4x), Saos-2 cell alkaline phosphatase level (2x) and mineralization (3x). The scaffolds are spongy in nature in a wet state, thus making them potential implants for bone cavities with a small opening. The application of these cryogenically 3D printed compressible gel scaffolds with multiscale porosity extends to small as well as a large sized open/partially open patient-specific bone defect.

Keywords: Cryogenic 3D Printing, Hydrogel, Bone Tissue Engineering, Multiscale Porosity, Gelatin, Carboxymethylchitin, Hydroxyapatite

1 Introduction Bone tissue engineering (BTE) is an emerging domain which aims to develop biological substitutes to provide a temporary framework for the regeneration of a functional tissue following loss due to traumatic disease and injury 1-2. These substitutes are intended to mimic the complex 3D micro-architecture of the natural bone tissue and its chemical and mechanical properties

3-4.

Highly porous 3D scaffold offers a great advantage of high surface area which is crucial for cell adhesion, proliferation, and migration 5. Further, interconnected open pores are essential for unimpaired diffusion of oxygen, nutrients, and wastes for proper vascularization and tissue growth 5-10.

Efforts have been made to incorporate micro as well as macroporosity into the scaffolds and

in some studies, it was found that scaffolds with multiscale porosity involving both micro and macro pores perform better in-vivo than only macroporous scaffolds 11-13. 3D printing emerged due to its potential of producing patient-specific scaffolds with controlled geometry by data obtained from magnetic resonance imaging (MRI) or computed tomography (CT) scans 4. For BTE, it has been explored to use numerous biomaterials and their composites via either melt extrusion, melt hybrid electrospinning-3D printing, cell encapsulated constructs and many others. It has shown immense potential in BTE as it overcomes the need of moulds which alleviates the difficulties in building complex shaped structures. Moreover, when 3D printing is accompanied by some other technology, additional functionalities can be added. For instance, Owen et al. developed PolyHIPE 3D hierarchical structures, where microstructure could be independently controlled by emulsion templating while scaffold macrostructure was controlled


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by additive manufacturing

14.

This combination of stereolithography and emulsion templating

could achieve interconnected multiscale porosity and including UV-light absorber reduced surface skin formation, improved resolution of stereolithography fabricated structures, and promoted osteoblast proliferation and bone matrix deposition

15.

Although such a combination of

technologies can generate multiscale porosity with photocurable polymers, it poses difficulties with water dissolvable natural polymers which are usually used to prepare hydrogels. These hydrogels often fail to withstand their own weight and lose their desired shape when multiple layers are formed during 3D printing 16. Deformation of the bottom layers and structure collapse can easily happen when printing low viscosity materials 17. This kind of gravity-driven distortion of a gel shape leads to the development of mismatched bone graft, especially for a customized bone defect. Adamkiewicz et al. introduced a novel cryogenic 3D printing method for 3D printing hydrogels on a cold plate by using liquid nitrogen as coolant

18.

In another study, solid carbon

dioxide was used to rapidly cool a composite hydrogel to create 3D geometrical structures and their mechanical properties were investigated

16.

The idea behind cryogenic 3D printing is to

transform a solution state gel, extruded from a nozzle, into a solid state thus allowing the building of stable frozen gel structures devoid of any gravity-driven deformation or collapse, without any need of support bath 16. However, lyophilization of the cryogenic 3D printed gel objects to create bulk micro-porosity and surface pores in the scaffolds has not been reported yet. A 3D artificial construct should be made of biomaterials which chemically mimics the natural bone. Natural bone is composed of 65 wt.% apatite crystals, 25 wt.% collagen fibers, and 10 wt.% water 7. Hence, materials used in this study for scaffold fabrication using 3D printing are composites of natural polymers namely gelatin (Gel), carboxymethyl-chitin (CMC) and ceramic hydroxyapatite (HA) [Ca10(PO4)6(OH)2]. Gelatin and CMC fulfill the requirement of organic content while HA serves as an inorganic framework 19. HA is a bioactive synthetic ceramic widely used to mimic apatite crystals due to its good biocompatibility, osteoconductivity, osteointegrative properties as well as its structural and chemical similarity with the inorganic component of the bone matrix 7-8, 20-25. Also, its better affinity and chemical bonding with the host hard tissues gives it an advantage over other bone substitutes

23.

However, its poor processability, brittleness,

fragility, hardness, lack of flexibility, high crystallinity and slower resorption rate in-vivo limits its application as a load bearing implant material 9, 26-28. To overcome these issues, HA is usually used



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as a composite with polymers not only to improve its processability and mechanical properties but also to enhance the osteoconductivity of the polymers 28. Polymers used for bone tissue engineering can be synthetic or natural. Synthetic polymers display controlled degradation rate and a potential to fabricate complex shapes. But one of their critical drawbacks is poor interaction with the host cells which limits their applications

29-30.

Natural polymers, on the other hand, show better bioactive properties and cellular interactions but exhibit poor mechanical strength and an uncontrolled degradation rate 31. Also, it still remains a challenge to generate macroporosity with natural polymers because of their solubility in aqueous media which confines their usage to a few fabrication techniques 19. The selection of biomaterials was based on the criterion that serves as a temporary extracellular matrix (ECM) for the regenerative cells. However, the process of neo-tissue genesis is not precisely the same as the developmental process. Hence, the scaffold was not aimed to entirely duplicate the natural ECM 32.

Since collagen is associated with pathogenic and immunogenic issues

33,

its biocompatible

derivative i.e. gelatin, which has been extensively utilized in orthopedics 34, was used. Gelatin has many advantages such as its biocompatibility, biodegradability, formability, non-toxicity, costeffectiveness, easy availability, lower antigenicity compared to collagen and cell recognition sites which improve final biological behavior and it also modulates cell adhesion 35-36. The presence of functional groups on the surface of gelatin can alter the surface charge thus enhancing cell attachment and proliferation 37. Further, some functional groups in the molecule like –NH2 and – COOH can be crosslinked which allows fabricating composite scaffolds which are stable in aqueous medium 35. Carboxymethyl chitin sodium salt (CMC), an amino polysaccharide, is one of the most usable derivatives of chitin mainly in the field of biomedical applications because of its unique biological and physiological properties like biocompatibility, non-toxicity, biodegradability, inexpensiveness and anti-microbial activity 38-41. In contrast to chitin which is poorly soluble in all common solvents 42

and chitosan which has limited solubility in water and other organic solvents, CMC, on the other

hand, is hygroscopic and completely soluble in water milder immune response

41.

43

and shows higher degradation rate and

It has been found that –COO– groups present in CMC not only

combine with the positive ions present in culture medium (in-vitro) and in the extracellular fluid (in-vivo) but also interact with Ca+2 ions of HA, which in turn mediates the positive cellular


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response and helps in proliferation and osteogenesis

44-45.

It also instigates a catalytic effect for

effective apatite formation on surface 46. Our objective in this study was based on a biomimetic approach, in which we used the above mentioned polymer-ceramic composites (Gel/CMC/HA) in an advanced 3D printing assembly, which can print gels with a wide range of polymer concentration. We compared four types of gel scaffolds differentiated on the basis of their manufacturing process. The influence of gel printing in cold environment followed by lyophilization on the generation of microporosity was analyzed. The effect of total polymer concentration on the printability was discussed. Further, the influence of different crosslinker and total polymer concentration on closed and open surface pores was studied. Finally, we successfully demonstrated the advantage of fabricating highly porous 3D printed constructs on cellular attachment, infiltration, growth and phosphatase generation and mineralization.

2 Materials and Methods 2.1 Materials RapidBot 3 was purchased from Makemendel, India and a block heater for holding a glass syringe was fitted in place of filament extruder. 30 mL glass syringe was procured from Top Syringe Mgf. Co. (P). Ltd., India. The SL101N digital dispensing controller was purchased from Fisnar Inc., USA. Carboxymethyl-chitin (CMC) was procured from Hangzhou Dayangchem Co. Ltd., China. Hydroxyapatite (HA, 50-60 nm nano-rods) in powder form was purchased from Plasma Biotal Ltd., UK. Gelatin type A (extracted from porcine skin), was provided by Rama Industries, India. 1-ethyl-3-(3-dimethylaminopropyl) carbodiimide hydrochloride (EDC) was purchased from Spectrochem Pvt. Ltd., India. For biological analysis, Saos-2 cell line was purchased from NCCS (Pune, India). McCoy’s 5A Medium cell culture media with L-glutamine (HiMedia, India), supplemented with streptomycin (100 µg/ml), penicillin (100 units/ml) and 10% fetal bovine serum (FBS) were procured from HiMedia, India. For hUCMSCs cell culture, Dulbecco’s Modified Eagle Medium (DMEM) was bought from Gibco, Invitrogen, USA which was supplemented with 2 mM L-glutamine, streptomycin (100 µg/ml), penicillin (100 units/ml) and 16% FBS. For cell proliferation assay, thiazolyl blue tetrazolium bromide (MTT) was purchased from Thermo Scientific Inc. (United States). Phosphate buffered saline (PBS) tablets were procured from HiMedia, India. SensoLyte® pNPP Alkaline Phosphatase Assay Kit was



purchased from AnaSpec, India. Quant-iT PicoGreen® DNA assay kit was purchased from Invitrogen, USA. Alizarin Red S was procured from HiMedia, India. 4′,6-diamidino-2phenylindole (DAPI), propidium iodide (PI) and FITC phalloidin were procured from Invitrogen, ThermoFischer Scientific, India. All other chemicals including trypsin-EDTA, Triton X-100, paraformaldehyde (PFA), glutaraldehyde, dimethyl-sulfoxide (DMSO) were purchased from Sigma Aldrich, India.

2.2 Fabrication of gel scaffolds The scaffolds were fabricated by 3D printing gel using an indigenously modified 3D printer as shown in Figure 1 which includes the customization of gel extrusion temperature as well as the build plate temperature. The composite gel solution was prepared by first suspending HA in MilliQ water in a glass bottle kept at 37 °C on a heated bed with an overhead stirrer. Gelatin was added and the stirring was continued until it was completely dissolved in water. CMC was then added slowly with continuous stirring and the solution was continuously stirred with an overhead stirrer for 12 h at 37 °C. After 12 h, a homogeneous solution was achieved which was then poured into a 30 mL glass syringe. The solution then cooled down to room temperature and turned into gel automatically. The syringe was mounted on the aluminum syringe-holder cum block heater which was attached with the 3D printer (Figure 1 A). The heater has a specifically cut hollow shape in which the syringe along with the needle can exactly fit at its center. Cartridge heaters were fitted to heat the syringe holder thus heating the gel uniformly. Heating was required to decrease the viscosity of high concentration gel which otherwise could not be extruded suitably. However, at a lower concentration, the gel did not require heating and could be printed at room temperature. After attaining the desired temperature, the air pressure was applied from the top to extrude the gel through a 22-gauge stainless steel needle on a cold surface (Figure 1 B). Each construct was made with the geometry of struts aligned with a repetition pattern of 0/0/90/90/… degrees. The cold plate assembly is shown in Figure 1 C. Briefly, an aluminum plate with vertical fins on one side was fixed in a closed hollow polystyrene box. Another thin aluminum plate (of 2 mm thickness, without fins), with the comparatively larger flat surface area, was attached on the top of the first plate. Liquid nitrogen was poured into the box through a hole carved at one of its edges. The fins of the aluminum plate immediately experienced a temperature gradient and thus started transferring heat from the outside, through the thin plate, into the box. This caused the thin


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plate to turn cold. A removable assembly was fitted which provided a continuous stream of cold nitrogen gas directly below the nozzle where the gel gets extruded.

Figure 1. (A) Gel 3D printing set up suitable for a wide range of polymer/ceramic concentrations. The aluminum block heater is specially designed to grip the syringe-needle system. The cartridge heaters heat it to lower the gel viscosity to get a smooth printing. (B) shows a uniform distribution of air pressure with a plunger for the gel to get extruded through a needle in a local cold environment which instantly freezes the extruded gel. For building largesized scaffolds, cold N2 air is blown over the printing area to freeze the upper layers. (C) shows the exploded view of the cold plate, where a finned plate is used for efficient heat transfer and an aluminum plate of a larger area is placed at the top. (D, E, F, G, H, I) show compressibility of the wet gel scaffolds. The scaffold shown in (D) has diameter 27±0.5 mm and height 10 mm. It is compressed till its width reaches 25% of its original value (F), and is recovered up to 90% of its original diameter after 1st cycle (H) and 85% after 3rd cycle (I).

A wide range of total polymer concentration (Gel and CMC) were tried for 3D printing with this set-up and it lies from 3% (Gel:CMC::1.5:1.5) to 20% (Gel:CMC::10:10) in water solution, and HA concentration was varied from 0 to 15% in water solution. In particular, we reported only two of them, first was Gel:CMC:HA::7.5:7.5:10 (i.e. 15% total polymer concentration and 10% ceramic in water), and second was Gel:CMC:HA::5:5:10 (i.e. 10% total polymer concentration and 10% ceramic in water). First set of scaffolds (Gel:CMC:HA::7.5:7.5:10 in solution) was 3D



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printed on a build plate kept at room temperature (25 °C) followed by drying at 25 °C (named as R15D, where ‘R’ stands for room temperature printing, ‘15’ for total polymer concentration i.e. Gel+CMC, and ‘D’ for drying), the second set had the same material composition as well as the printing temperature but followed by lyophilization at -50 °C and 50 hPa instead of drying (named as

R15L,

where

‘L’

stands

for

lyophilization),

and

the

remaining

two,

with

Gel:CMC:HA::7.5:7.5:10 and Gel:CMC:HA::5:5:10, were printed in a cold atmosphere followed by lyophilization and named as C15L and C10L respectively, where ‘C’ stands for printing in a cold atmosphere. All the scaffolds were crosslinked by dipping them into EDC solution of predefined concentration followed by thorough rinsing with Milli-Q water to remove excess EDC. R15L, C15L, and C10L were then again lyophilized while R15D was dried at room temperature. The crosslinking happens between carboxyl and amine groups of CMC and Gel. After completion of the fabrication procedure, the dried R15D, R15L, and C15L scaffolds had Gelatin:CMC:HA ratio of 30:30:40, while C10L had the ratio of 25:25:50. The nomenclature of scaffolds with their fabrication conditions is tabulated in Table 1. Table 1 Fabrication conditions of the scaffolds. The scaffolds nomenclature can be explained as, R: printing at room temperature, C: printing in cold atmosphere, D: drying at 25°C, L: lyophilization, 15 and 10: total polymer concentration of 15% and 10% in solution respectively.

Scaffold

Total polymer concentration in solution (mg/ml)

Build plate temperature (°C)

Gel:CMC:HA ratio in solution (%)

Drying process

R15D

15

25

Normal drying

7.5:7.5:10

30:30:40

R15L

15

25

Lyophilization

7.5:7.5:10

30:30:40

C15L

15

-72 ± 8

Lyophilization

7.5:7.5:10

30:30:40

C10L

10

-72 ±8

Lyophilization

5:5:10

25:25:50

Gel:CMC:HA ratio in scaffold (%)

2.3 Physical characterization 2.3.1

Morphological analysis The morphological analysis of the samples was analyzed using field emission gun scanning

electron microscopy (SEM, JEOL JSM- 7600F) equipped with energy dispersive X-ray spectroscopy (EDS). The samples were first mounted on the stubs and coated with a thin platinum layer for 300 seconds using sputter coater (SC7640 Sputter Coater, Quorum Technologies Ltd,


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UK) to make the polymer surface conductive. The samples were then analyzed under SEM at an accelerating voltage of 15 kV. 2.3.2

Fourier Transform Infrared Microscopy (FTIR) FTIR spectra were obtained using Nicolet Magna-IR FTIR 550 spectrophotometer (Nicolet

Instrument Corporation, Madison, WI, USA). The spectral range was kept from 4000 to 400 cm−1. 2.3.3

Porosity calculation Scaffold porosity was calculated by using liquid displacement technique similar to that

reported in the literature 47. Ethanol was used for porosity calculation. As it is a non-solvent for the polymers, it does not induce swelling or shrinkage. Scaffold sample was introduced in a cylinder having a known volume of ethanol (V1). After 5 min of immersion followed by multiple brief evacuations and repressurization cycles for forced ethanol penetration into the pores of the scaffold. The cycle was continued until there was no bubble found emerging from the scaffold. The volume of ethanol containing scaffold was noted as V2. The difference in volume (V2 – V1), gives the total volume of the scaffold material. The ethanol impregnated scaffold was removed from the cylinder and the residual ethanol volume was recorded as V3. The volume difference V1 – V3 gives the void volume of the scaffold. Another similar experiment was performed with scaffold R15D having no porosity and the residual ethanol volume was recorded as V3’. The volume difference V1 – V3’ in this case gives the measure of ethanol uptake and/or absorbed by the material and is termed as Vabs. Hence, the effective void volume of any porous scaffold was calculated from V1 – V3 – Vabs. The total volume of the scaffold is given by V = (V2 – V1) + (V1 – V3 – Vabs) = V2 – V3 – Vabs The percentage bulk porosity of the scaffold (ε) is then given by the following formula (Eq. 1): ε = [(V1 ― V3 ― Vabs)/(V2 ― V3 ― Vabs)] ∗ 100 2.3.4

Eq…1

Water uptake The swelling property or water uptake ability of the scaffolds was evaluated by immersing

the pre-weighed dried scaffolds in Milli-Q water at 37°C. Post incubation, the scaffolds were taken out after predefined time intervals and excess water was removed using tissue paper and the wet scaffolds (Ww) were weighed. The scaffolds were then dried in an oven to their constant weight (Wd). The water uptake was calculated using the following formula (Eq. 2):



𝑆𝑤𝑒𝑙𝑙𝑖𝑛𝑔 (%) = 2.3.5

𝑊𝑤 ― 𝑊𝑑 𝑊𝑑

∗ 100

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Eq…2

Degradation in SBF The degradation profile of the scaffolds was measured by immersing them in a simulated

bodily fluid (SBF). The scaffolds were placed in 24 well plates with 2 mL of SBF in each well. SBF was replenished after every 3 days. The scaffolds were taken out after predefined time points and washed gently with Milli-Q water and dried to constant weight. The % degradation was calculated by using the following formula (Eq. 3): %∆𝑊 =

𝑊𝑓 ― 𝑊𝑖 𝑊𝑖

∗ 100

Eq…3

Where, Wi is the initial weight, and Wf is the final weight. 2.3.6

Compression testing Uniaxial compression testing of dry scaffolds in z-axis was carried out using a Universal

Testing Machine with a load cell of 5 kN and a compression rate of 0.1 mm/s. Cylindrical scaffolds of diameter 10 mm and height 10 mm were compressed till 70% compression or till complete compaction occurred, whichever was earlier.

2.4 Biological assessment of gel scaffolds This study provides an insight into the biocompatibility of gel scaffold in terms of attachment, proliferation, protein generation and mineralization of osteosarcoma Saos-2 cell line. Saos-2 cell line was selected as it has been considered as a useful in-vitro model for studying human osteocyte differentiation and production of bone-like mineralised matrix. Additionally, the cyto-compatibility of these scaffolds has also been tested with primary cells (mesenchymal stem cells) which have normal cell morphology and maintain many crucial markers and functions as seen in-vivo which a cancerous cell line lacks. 2.4.1

Study of human umbilical cord mesenchymal stem cell (hUCMSCs) attachment and proliferation on gel scaffolds

2.4.1.1 Isolation of hUCMSCs from umbilical cord tissue Umbilical cords (UCs) were collected in a 50 mL centrifuge tube filled with phosphate buffered saline (PBS) supplemented with antibiotics (100 µg streptomycin and 100 U/mL penicillin) from a local maternity and health care after signing the informed consents. The protocol


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for the isolation of hUCMSCs from UC was adapted from Verma et al. 48. Briefly, the cord blood was squeezed out using curved forceps and washed with sterile PBS. Blood from the vessels was flushed with sterile PBS. After cleaning, the UC was processed for the isolation of hUCMSCs. Firstly, the blood vessels were removed by blunt dissection. The remaining cord was chopped into pieces of length 1-2 mm using a sterile surgical blade. The chopped tissues were subjected to digestion step in an enzymatic cocktail made of collagenase type IV and Dispase of ratio 7:1 v/v at 37°C on a magnetic stirrer. After 30 min, trypsin (0.05 wt.%) supplemented with EDTA (0.02 wt.%) was added and the mixture was again stirred at 37°C for 15-20 min. The formed homogenate was then filtered using a sterile nylon filter. The filtrate was centrifuged at 239 g for 10 min to obtain a cell pellet. The supernatant was discarded and the pellet was re-suspended and cultured in DMEM and incubated at 37°C and 5% CO2. After extraction of MSCs from the umbilical cord, the mixed population of cells was purified using FACS (BD FACS Aria Special Order System) with MSCs marker CD44. The pure MSCs population obtained from FACS was then propagated to achieve the desired cell number for experiments. The culture media was changed after every two days. 2.4.1.2 Characterization of isolated hUCMSCs The isolated hUCMSCs were dislodged using 0.05% trypsin (supplemented with 0.02 % EDTA) at the 3rd passage and seeded on a coverslip. The cells were then fixed using 4% paraformaldehyde at 4°C for 15 min. It was followed by cell membrane permeabilization with 0.2% Triton X-100. After blocking the cells with bovine serum (HiMedia, India), they were washed gently. The cells were then incubated in primary antibodies, FITC mouse anti-human CD90 (Biolegend), and anti-Oct-4 (Biolegend) for 2 h. After giving gentle PBS wash, a secondary antibody Alexa Fluor 488 goat anti-mouse IgG (Invitrogen, CA, USA), was added. The fluorescence images were taken using spinning disk confocal microscopy (Zeiss, Germany). 2.4.1.3 hUCMSCs cell seeding Gel scaffolds were first sterilized by dipping them in 70% ethanol followed by 30 min UV light exposure. The excess ethanol was removed by thorough washing with sterile PBS. The scaffolds were preconditioned in DMEM culture media for 2 h. 5x104 cells were seeded to each scaffold kept in a 24 well-plate. The cells were allowed to attach for 4 h. After that, the scaffolds



with cells were transferred to another 24 well-plate and fresh DMEM media was added. The culture media was replenished every 2 days. 2.4.1.4 Viability assay The cell viability on the scaffolds was measured using MTT (3-(4,5-Dimethylthiazol-2yl)-2,5-diphenyltetrazolium bromide). 1 mg/mL of MTT (Sigma Aldrich, USA) was prepared in PBS and filtered using a 0.2 µm filter. The media was removed from the scaffolds and washed with PBS. 200 µL of MTT solution was added to each scaffold which was then incubated at 37°C for 4 h. The reducing agents present in metabolically active cells reduced the yellow MTT to violetblue formazan. 800 µL of dimethyl sulfoxide was added to each scaffold to dissolve the formazan crystals. The optical density of the formazan crystals was measured at 560 nm using ELISA plate reader (Thermo-Fischer Scientific Inc., United States). Control scaffold without cells was tested and the values were subtracted from all the groups and the subtracted values are reported here. The same protocol was followed for the assessment of Saos-2 cellular viability. 2.4.1.5 Proliferation assessment Cellular proliferation of hUCMSCs and Saos-2 cell lines were assessed quantitatively with DNA quantification assay with Quant-iT PicoGreen® DNA assay kit. The cell-seeded scaffolds were harvested at predefined time points. The scaffolds were washed three times with PBS and then dipped in deionized water for cell lysis. Thereafter, the scaffolds were frozen and thawed alternatively two times followed by centrifugation at 10,621 g for 10 min. Then 100 µl of PicoGreen® working solution was added to the supernatant and incubated for 5 min. The DNA concentration was measured using a microplate fluorescence reader (Spectramax Molecular Devices, USA) at an excitation of 480 nm and emission of 538 nm. Finally, the DNA content was measured by using the standard curve prescribed by the manufacturer’s kit. The same protocol was followed for the assessment of Saos-2 cellular proliferation. 2.4.2

Saos-2 cell culture study on gel scaffolds

2.4.2.1 Cell culture conditions The human osteosarcoma Saos-2 cell line was cultured in McCoy’s 5A Medium cell culture media with L-glutamine, supplemented with streptomycin (100 µg/ml), penicillin (100 units/ml) and 10% fetal bovine serum and maintained in a humidified incubator kept at 37°C with 5% carbon dioxide (CO2). For the assays, the prepared gel scaffolds were placed in 24 well-plates (Eppendorf,


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USA) and were sterilized with 70% ethanol and UV exposure for 30 min each. The scaffolds were dried and preconditioned by incubating them in cell culture media for 2 h. 5 x 104 cells were seeded on the scaffolds followed by 4 h incubation at 37°C under 5% CO2 for the cells to attach on scaffold surface. Afterward, 1 mL media was added in all the wells and was replenished every two days. 2.4.2.2 SEM analysis for attachment studies The cell-seeded samples were washed with PBS and the cells were fixed with 2.5% glutaraldehyde for 2 h. The excess glutaraldehyde was removed with PBS wash and the samples were dried with sequential ethanol treatment, from 25% ethanol to 100% ethanol and then kept at 37°C overnight for drying. The samples were then mounted on stubs and coated with platinum and the SEM images were taken at an accelerating voltage of 15 kV. 2.4.2.3 Confocal studies The media was removed after 7 days and the scaffolds were washed with PBS. 4% paraformaldehyde was then added for 2 h to fix the cells. The scaffolds were washed with PBS after removing the paraformaldehyde. After that, the cell membranes were permeabilized with 0.2% Triton-X 100 in order to facilitate the permeation of propidium iodide (PI) and FITC phalloidin into the cells. For staining, the cells were dipped in 2 units/mL FITC phalloidin stain at 4°C for 6 h to stain the actin filaments. It was followed by three gentle PBS wash of 2 min each to remove extra FITC from the scaffolds. The scaffolds were then dipped in 10 µg/mL DAPI stain or 20 µg/mL PI stain at 4°C for 15 min to stain the nucleus. The scaffolds were again washed with PBS three times and then analyzed under a laser scanning multiphoton confocal microscope (Carl Zeiss Meditec AG, Jena, Germany). The images were processed using Zen software (Zeiss). 2.4.2.4 Alkaline phosphatase (ALP) activity The ALP activity of Saos-2 cells was measured after 3, 7 and 14 days. The media was removed from the wells and the scaffolds were washed with PBS. 10X assay buffer (Component B) was first diluted to 1X with deionized water (Sensolyte, Anaspec). Then, 20 µL of Triton X100 (Component D) was added to 10 mL of 1X assay buffer to prepare the cell lysis buffer. The cells were harvested using the above-prepared cell lysis buffer. The harvested cell lysate was centrifuged at 2500 g for 10 min. The supernatant was then taken to measure the ALP activity using ALP activity kit according to the manufacturer’s instructions (Sensolyte, Anaspec). Briefly, the para-nitrophenyl phosphate (PNPP) (50 µL) was mixed with 50 µL of cell lysate suspension



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and incubated at 37°C for 60 min. After 60 min of incubation, the enzymatic reaction was stopped by adding 100 µL of 1 N NaOH solution and the para-nitrophenol production was observed by measuring the absorbance at 405 nm using ELISA plate reader. Finally, the ALP activity was determined from a standard curve. The total protein content was measured by using the BCA Protein Assay Kit with bovine serum albumin as a standard. 2.4.2.5 Alizarin red S staining (ARS) assay To quantify the cell mineralization, alizarin red S assay was carried out. The Saos-2 cells were fixed with 2.5% glutaraldehyde for 4 h after 3, 7 and 14 days of culture in scaffolds. The scaffolds were then washed thoroughly with double distilled water (DDW) to remove the excess glutaraldehyde. Then, 2 mM of ARS solution was added to the scaffolds and incubated for 20 min with shaking at room temperature to facilitate staining. Extra dye was removed by washing the cell-scaffold constructs with DDW thrice. For the quantification of calcium formed by the cells, 400 µL of 10% acetic acid was added to each well. After incubating the scaffolds in acetic acid for 30 min, they were transferred to a microcentrifuge tube and kept at 80°C for 10 min, cooled then centrifuged at 10,000g. 500 µL of the above-taken supernatant was neutralized with 10% ammonium hydroxide solution and the absorbance was measured at 405 nm. Control scaffolds without cells were tested and the values were subtracted from all the groups and the subtracted values are reported here. 2.4.3

Hemocompatibility of scaffolds Human blood was collected from a healthy blood donor by following the extant

institutional guidelines. The collected blood sample was first centrifuged at 166 g at 10°C for 15 min. The supernatant mainly containing platelet-rich plasma was removed. The pellet formed at the bottom containing erythrocytes was washed thrice with a normal saline solution (NSS) (0.9% w/v NaCl). The hematocrit then obtained was incubated with gel scaffolds at 37°C and 5% CO2 for 1 h in an incubator (Thermo ScientificTM, United States). The samples were centrifuged at 1000 g for 5 min. The quantification of RBC lysis was done at an optical density of 542 nm using UV– vis spectrophotometer (Molecular Devices, United States). NSS and DI water were used as negative and positive control respectively. The hemolysis ratio was calculated using the following formula (Eq. 4): 𝐴𝑆 ― 𝐴𝑁

𝐻𝑅 = 𝐴𝑃 ― 𝐴𝑁 ∗ 100


Eq…4

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where, AS, AP and AN stands for the absorbance of sample supernatant, positive control, and negative control respectively.

2.5 Statistical analysis All statistical analysis was done in Origin 2018. The data was analysed by one-way analysis of variance (ANOVA). Tukey’s test was used for calculating significant differences. The differences were considered significant for p

Multiscale porosity in compressible cryogenically 3D printed gel for

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