3D-Printed High Strength Bioactive ... - ACS Publications

Subscriber access provided by UB + Fachbibliothek Chemie | (FU-Bibliothekssystem)

Article

3D-printed high strength bioactive supramolecular polymer/ clay nanocomposite hydrogel scaffold for bone regeneration Xinyun Zhai, Yufei Ma, Chunyong Hou, Fei Gao, Yinyu Zhang, Changshun Ruan, Haobo Pan, W. Lu, and Wenguang Liu ACS Biomater. Sci. Eng., Just Accepted Manuscript • Publication Date (Web): 17 May 2017 Downloaded from http://pubs.acs.org on May 18, 2017

Just Accepted “Just Accepted” manuscripts have been peer-reviewed and accepted for publication. They are posted online prior to technical editing, formatting for publication and author proofing. The American Chemical Society provides “Just Accepted” as a free service to the research community to expedite the dissemination of scientific material as soon as possible after acceptance. “Just Accepted” manuscripts appear in full in PDF format accompanied by an HTML abstract. “Just Accepted” manuscripts have been fully peer reviewed, but should not be considered the official version of record. They are accessible to all readers and citable by the Digital Object Identifier (DOI®). “Just Accepted” is an optional service offered to authors. Therefore, the “Just Accepted” Web site may not include all articles that will be published in the journal. After a manuscript is technically edited and formatted, it will be removed from the “Just Accepted” Web site and published as an ASAP article. Note that technical editing may introduce minor changes to the manuscript text and/or graphics which could affect content, and all legal disclaimers and ethical guidelines that apply to the journal pertain. ACS cannot be held responsible for errors or consequences arising from the use of information contained in these “Just Accepted” manuscripts.

ACS Biomaterials Science & Engineering is published by the American Chemical Society. 1155 Sixteenth Street N.W., Washington, DC 20036 Published by American Chemical Society. Copyright © American Chemical Society. However, no copyright claim is made to original U.S. Government works, or works produced by employees of any Commonwealth realm Crown government in the course of their duties.

Page 1 of 37

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

ACS Biomaterials Science & Engineering

3D-printed high strength bioactive supramolecular polymer/clay nanocomposite hydrogel scaffold for bone regeneration Xinyun Zhai,†,‡,§ Yufei Ma,‡ Chunyong Hou,‡ Fei Gao,† Yinyu Zhang,† Changshun Ruan,*,‡ Haobo Pan,‡ William Weijia Lu,*,‡,§ and Wenguang Liu*,† †

School of Materials Science and Engineering, Tianjin Key Laboratory of Composite

and Functional Materials, Tianjin University, Tianjin 300352, China. ‡

Research Center for Human Tissue and Organs Degeneration, Institute Biomedical

and Biotechnology, Shenzhen Institutes of Advanced Technology, Chinese Academy of Sciences, Shenzhen 518055, China. §

Department of Orthopaedic and Traumatology, The University of Hong Kong, 21

Sassoon Road, Pokfulam, Hong Kong, China.

Keywords: 3D-Printing, high strength, nancomposite hydrogel scaffold, bioactive, bone regeneration

ABSTRACT: The emerging 3D bioprinting technique that is strongly dependent on the development of bioinks offers a promising opportunity to customize personalized bioscaffold for precision and individualized therapy of bone defects. Hydrogels are one sort of attractive scaffolding materials due to their resemblance to extracellular matrices. Although much progress has been made in designing and fabricating high strength hydrogels, very few of them have been extended to the treatment of bone 1

ACS Paragon Plus Environment


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

defects. In this work, we developed a hybrid bioink composed of hydrogen bonding monomer (N-acryloyl glycinamide) (NAGA) and nanoclay. The hybrid ink could be conveniently tailored as a high strength PNAGA-Clay composite scaffold under UV light illumination of printed pre-hydrogel. The hydrogen bonding combined with physical crosslinking of nanoclay contributed to the superior mechanical performances as well as swelling stability of the hydrogels and bioscaffols. The sustainable release of intrinsic Mg2+ and Si4+ from the PNAGA-Clay scaffold was shown to promote the osteogenic differentiation of primary rat osteoblasts (ROBs) cells. Importantly, this implantable PNAGA-Clay scaffold highly efficiently facilitated the regeneration of new bone in tibia defect of rats. We anticipate hybridization of hydrogen bonding monomer with a variety of bioactive inorganic nanoparticles will offer new possibilities to develop numerous bioinks for 3D printing of desired bioscaffolds to realize individualized repair of degenerated load-bearing tissues. 1. INTRODUCTION Bone defects occur in a wide variety of situations such as trauma, developmental deformities or tumor resection. Reconstruction of bone defect with suitable scaffolds finely fitting the original surrounding bone tissues is one of important options for patients’ rehabilitation. Various degradable or nondegradable scaffold materials have been developed for bone tissue regeneration by providing three-dimensional architecture emulating local microenvironment.1-6 Among the optional material candidates, hydrogels have numerous attractive features as bone scaffolds due to their 2


Page 2 of 37

Page 3 of 37

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


resemblance of extracellular matrix as well as a broad range of modulus tunability.7-12 However, the poor mechanical properties intrinsic to conventional hydrogels severely restrain their application as loading support for bone regeneration.13-14 Past decade witnessed the rapid growth of high strength hydrogels mostly from the theoretical point of view or proof of concept.15-21 Most recently, hydroxyapatite (HAp) nanocrystals coated double-network tough hydrogels were reported, and the designed plug was implanted into an osteochondral defect created in the rabbit femoral groove. This hybrid construct exhibited strong binding to bones due to the spontaneous osteogenesis penetration.22 Our team also investigated on a mineralized dipole-dipole reinforced high strength and tough hydrogel for skull bone regeneration.23 Nonetheless, these chemically crosslinked high strength hydrogels lost design and reprocessability for fabricating complex geometric scaffolds, let alone the potential toxicity of crosslinker used. The emerging 3D bioprinting technique can allow for customizing personalized tissue engineering scaffold for precision and individualized therapy.24-25 The successful application of 3D printing in engineering tissue scaffolds is strongly dependent on the development of bioinks.26-27 As an ideal hydrogel-based bioink for printing load-bearing scaffold, it ought to satisfy these rigorous criteria: (1) the monomer solution should be extrusible and 3D-printable; (2) the extruded solution should be gelled quickly enough on the printing substrate to self-support layer-by-layer construct; (3) the multilayer stacks should be bound tightly to prevent delamination; (4) the printed hydrogel-based scaffolds to bear loading should be 3



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

stable in physiological aqueous condition without further swelling. We previously invented a supramolecular polymer hydrogel, poly(N-acryloyl glycinamide) (PNAGA) whose dual amide motif in side chain could form stable hydrogen bonded crosslinking domains, thus contributing to the significant increase in mechanical strengths and swelling stability.28 However, this high strength inert hydrogel is devoid of bioactivity. It has been recognized that nanoclay (Laponite XLG) is a biocompatible and bioactive inorganic multifunctional crosslinker which has been successfully used to construct high performance hydrogels due to the uniformly physical crosslinking of polymer chains by neighboring clay sheets.29-30 Recently, Zhao et al used Laponite clay as an additive to prepare PEG-alginate-clay bioink which was then printed into complex cellularized structures incorporating human embryonic kidney cells.31 This is the first work demonstrating that highly stretchable and tough hydrogel was printable and suitable for long-term cell culture. However, no research has been done to extend 3D-printed high strength hydrogels to genuine bone regeneration thus far. In this study, we propose to print 3D high strength supramolecular polymer/clay nanocomposite hydrogel scaffold. We anticipate that nanoclay could finely tune the viscosity of hybrid bioink to prevent the printed multilayer scaffold from collapsing on the printing substrate, and dual amide hydrogen bonding interactions combined with nanoclay crosslinkage would eventually result in high strength and swelling stable hydrogels, thus providing reliable loading support. In addition, the release of bioactive ions could stimulate the osteogenesis in vivo. 2. EXPERIMENTAL SECTIONS 4


Page 4 of 37

Page 5 of 37

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


2.1. Materials. Glycinamide hydrochloride (98%, TCI, Shanghai, China), acryloyl chloride (98%, TCI, Shanghai, China), diethyl ether (Lingfeng Chemical Reagent Company, Shanghai, China), 2-hydroxy-2-methyl-1-phenyl-1-propanone (IRGACURE 1173, 98%, Sigma-Aldrich, St. Louis, USA) and Laponite XLG ([Mg5.34Li0.66Si8O20(OH)4]Na0.66) (BYK, Wesel, Germany) were used as received. All other chemicals and solvent are analytical reagents. 2.2.

Preparation of pre-gel solution for 3D-printing. NAGA was synthesized as

described in our previous work.28 NAGA monomer was firstly dissolved in deionized water according to the specified recipes, and then 3 wt% of photoinitiator IRGACURE 1173 (relative to the mass of NAGA monomer) was added into the solution and stirred thoroughly under nitrogen atmosphere until completely dissolved. Subsequently, nanoclay (Laponite XLG) was added into the monomer solution and vigorously stirred in an ice-water bath for about 2h under nitrogen protection. Bubbles were removed from the pre-gel solution by centrifugation before used. In this experiment, 10%, 20% and 30% NAGA contents (in mass percentage, relative to water) were used to prepare hydrogels. 2.3.

3D-Printing of scaffolds and hydrogel formation. The printing device (3D

scaffold printer) is a precision three-axis positioning system (BioScaffolder 2.1, GeSiM, Grosserkmannsdorf, Germany). The dosing pressure to the syringe pump was controlled between 25 kPa to 35 kPa and the moving speed of the dispensing unit was set to 15 mm s-1. Different plotting parameters and nozzle sizes were selected to control the morphology of the scaffolds. For the in vitro and in vivo study in our 5



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

experiments, nozzle size was chosen to be 250 µm, the distance between each strand was about 1 mm and the scaffolds with 10 layers were fabricated. The above pre-gel solution was transferred into the printer cartridges. The thick solution was readily extruded out of needle due to shear thinning, and quickly recovered its viscosity upon dropping down on the printing substrate. Immediately, the printed pre-scaffolds were placed into a crosslink oven (XL-1000 UV Crosslinker, Spectronics Corporation, NY, USA), and the polymerization was performed for 30 min at 365 nm. The obtained scaffolds were washed thoroughly with distilled and deionized water to remove the impurities, and stored in water at room temperature. The resulting scaffolds were named as PNAGAX-Clay (X is NAGA concentration in mass percentage, relative to water). PNAGAX hydrogels and PNAGAX-Clay hydrogels were also prepared by UV light-initiated polymerization without using 3D printing. 2.4.

Basic characterizations of PNAGA-Clay. Fourier transform infrared

spectroscopy (FTIR) (BRUKER VERTEX 70, Madison, USA) was used to characterize the dry scaffold samples. Thermogravimetric analysis (TGA) (Q600 SDT, TA Instruments, New Castle, USA) was used to characterize pure clay, PNAGA and PNAGA-Clay scaffold.32 The sample was heated from room temperature to 1090 °C at heating rate of 10 °C/min in air. For transmission electron microscopy (TEM) imaging,33 lyophilized PNAGA20%-Clay scaffolds were embedded in epoxy resin for microtoming at −40 °C with a glass knife to ca. 50 nm thick sections, which were collected onto copper grids for imaging in a transmission electron microscope (Philips 6


Page 6 of 37

Page 7 of 37

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


CM100, TEM, Massachusetts, USA) at 100 kV later. The microstructures of freeze-dried PNAGA-Clay hydrogels were observed by a field emission scanning electron microscope (SEM, FEI Quanta S-4800 FE-SEM, Oregon, USA) at accelerated electron energy of 5.0 kV after gold sputtering. 2.5.

Mechanical test. All mechanical properties of the pure hydrogels and

PNAGA-Clay scaffolds were tested on the Instron 5697 (Instron, Grove City, USA) universal material testing system at room temperature. Before test, all the hydrogels and scaffolds were fully equilibrated in deionized water. For tensile test, PNAGA and PNAGA-Clay hydrogels were cut into rectangle with thickness in 1 mm, 7 mm in width and 50 mm in length. The extension rate was controlled at 50 mm min-1. The cylinder-shaped hydrogels (4 mm in diameter and 4 mm in height) and printed rectangle scaffolds with size of 4 × 4 × 4 mm were used respectively for compression test with the crosshead speed at 10 mm min-1. In order to get mean and standard deviation calculation, at least five samples were used for each test. 2.6.

Ion leaching analysis. The printed scaffolds with 10 layers (15 mm in width)

were immersed into 10 mL of phosphate buffer solution (PBS). At different immersion time points (1, 2, 3, 4, 5, 6, 7, 14 and 21 days), 50 µL solution was taken out and appropriately diluted. Then the diluted solution was analyzed by inductively-coupled plasma optical emission spectrometry (Perkin Elmer, Optima 7000DV, Massachusetts, USA) to determine magnesium ions and silicon ions concentrations released. At each time point, corresponding 50 µL fresh PBS was supplemented. 7



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

2.7.

Rheology test. Dynamic rheological experiments were carried out using the

rheometer (MCR302, Anton Paar, Austria). Plate-plate geometry with a plate diameter of 25 mm was used. Data for shear rates in the range 1-100 s-1 at room temperature were collected. 2.8.

Cell culture. Primary rat osteoblasts (ROBs) cells were isolated from minced

Sprague-Dawley rat (SD rat, born within 3 days) calvarial chips as described previously,34-35 and this procedure was approved by the Ethics Committee for Animal Research, Shenzhen Institutes of Advanced Technology, Chinese Academy of Sciences. ROBs cells were used to evaluate cell proliferation, cell viability and differentiation in vitro. The cells were cultured in alpha minimum essential medium (α-MEM) (Hyclone, Utah, USA) with 10% (v/v) fetal bovine serum (Corning, New York, USA) and supplemented with antibiotics (100U/mL penicillin, 100 µg/mL streptomycin) at 37 °C with 5% carbon dioxide (CO2). 2.9.

Proliferation assay and cytotoxicity assay. The ROBs cells were seeded

onto the sterilized scaffolds (7.5 × 7.5 × 2-3 mm) placed in 24-well cell culture plates (Corning, New York, USA) at a seeding density 5 × 104 cells per scaffold, and cultured as mentioned above. Cell proliferation was determined after 1, 3, 5 and 7 days by CCK-8 assay.36 The scaffolds were taken out from the medium and placed in a new 24-well cell culture plates. After incubation with 10% CCK-8 at 37 °C for 4 h, the number of viable cells was quantified by measuring the optical density of the CCK-8 solution at 450 nm using a Multiskan spectrum reader (Bio Tek Synergy4, Winooski, USA). The experiment was performed in triplicate. 8


Page 8 of 37

Page 9 of 37

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


The live-dead viability assay was performed according to the instructions to test the cytotoxicity of the ROBs cells seeded on the scaffolds after culturing for 7 days. The samples were observed using the confocal laser scanning microscope (CLSM) (Leica SD AF, Hamburg, Germany). The excitation filter was set at 488 nm to observe living (green) cells detected by calcein AM and at 561 nm to detect dead (red) cells detected by EthD-1.37 For cell morphology observation after seeded within the scaffolds, the cytoskeleton of ROBs cells was visualized with 50 µg mL-1 phalloidin-rhodamine (Cytoskeleton, Inc., Denver, USA) after fixation with 4% paraformaldehyde (PFA) and permeation by 0.1% Triton.38 And the growth status of cells after seeded within the scaffolds was observed by SEM (Hitachi S4800 FEG, Tokyo, Japan) after fixation, gradient dehydration, critical point drying and gold sputtering.38 2.10. Differentiation assay. The osteoblast phenotype of ROBs cells grown within the scaffolds was evaluated by quantification of alkaline phosphatase (ALP) activity and ALP staining.39 Cells are seeded on the scaffolds as described above. After 4, 7, 14 and 21 days’ culture, the scaffolds seeded with ROBs cells were rinsed three times with PBS and then lysed in 300 µL lysis buffer (RIPA buffer, Beyotime Biotechnology, Shanghai, China) for 30 min at 4 ºC. After centrifugation for 5 min at 13,000 rpm and 4 °C to remove cell debris, 50 µL of the supernatant was added to 50 µL chromogenic substrate and incubated at 37 °C for 2 h. The reaction was stopped by the addition of 100 µL stop buffer. Absorbance was measured at 405 nm using a microplate reader. Analysis of each sample was performed in triplicate and ALP activity was normalized 9



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

to the total protein content of each scaffold associated with the cell surfaces and matrix using a commercially available protein assay kit (PierceTM BCA Protein Assay Kit, ThermoFisher Scientific, Massachusetts, USA). For intuitive observation, ALP staining method was used by using BCIP/NBT Alkaline Phosphatase Color Development Kit (Beyotime Biotechnolgy, Shanghai, China). 2.11. Real-time quantitative RT-PCR analysis. The osteogenic differentiation properties of the PNAGA-Clay nanocomposite hydrogel were further assessed by real-time quantitative RT-PCR.40 The relative mRNA expression levels of three commonly used bone markers including alkaline phosphatase (ALP), type І collagen and osteocalcin (OCN) (primer pairs used are shown in Table 1) were measured. Since PNAGA hydrogel could not be printed into 3D-scaffold, all the sample disks were in 2-dimension with 5 mm in diameter and 1 mm in thickness. The ROBs cells were seeded on the sample disks placed in 24-well cell culture plates (Corning) at a seeding density 5 × 104 cells per sample and cultured as mentioned in cell culture part. After 7 days’ culturing, the total RNA of the osteoblast was extracted using TRIZOL reagent (Invitrogen, California, USA) and the complementary DNA (cDNA) was reverse-transcribed from 1 µg of total RNA using a RevertAid First Strand cDNA Synthesis Kit (ThermoFisher Scientific, Massachusetts, USA) according to the manufacturer’s instruction. The real-time PCR was performed on the SYBR Green PCR Master Mix (Toyobo Life Science, Osaka, Japan) and the reaction was carried out on the ABI 7500 (Applied Biosystems, Massachusetts, USA). Finally, the 10


Page 10 of 37

Page 11 of 37

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


house-keeping gene glyceraldehyde-3-phosphate dehydrogenase (GAPDH) was used to normalize the relative mRNA expression level of each gene and the quantification was based on the CT (cycle threshold) values. Table 1. Primer pairs used in real-time quantitative PCR analysis Gene

Forward primer (5’-3’)

Reverse primer (5’-3’)

Accession No.

GAPDH

GGCACAGTCAAGGCTGAGAATG

ATGGTGGTGAAGACGCCAGTA

NM_017008.4

ALP

AACGTGGCCAAGAACATCATCA

TGTCCATCTCCAGCCGTGTC

NM_013059

Collagen

GCCTCCCAGAACATCACCTA

GCAGGGACTTCTTGAGGTTG

NM_053304.

OCN

GGTGCAGACCTAGCAGACACCA

AGGTAGCGCCGGAGTCTATTCA

J04500

2.12. Surgical procedures for in vivo evaluation. All the animal procedures and experiments were approved by the Ethics Committee for Animal Research, Shenzhen Institutes of Advanced Technology, Chinese Academy of Sciences. A rat tibia model was used and the surgical procedures were conducted under sterile conditions. The scaffolds were cut into sections with 2.5 mm in diameter and 5 mm in length and were implanted into tibia bone defects of 12-week-old male Sprague-Dawley rats (SD rats). Briefly, the rats were anesthetized by intraperitoneal injection of pentobarbital sodium (40 mg/kg). Critical size defects (2.5 mm diameter and 5 mm length) were created in the medial of the tibia shaft, close to the tibia plateau for each SD rats. The defects were prepared with a 2.5 mm drill. The implants were placed bilaterally resulting in two implants per rat, and the wound was closed carefully. Eight SD rats were randomly divided into 2 groups corresponding to PNAGA20%-Clay scaffolds and blank as a control. 2.13. Sequential fluorescent labeling. In order to characterize the new bone 11



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

formation and mineralization, a polychrome sequential fluorescent labeling method was used.41 At 2, 4 and 6 weeks after operation, different fluorochromes were administered intramuscularly at a sequence of 25 mg/kg tetracycline hydrochloride, 30 mL/kg alizarin red S and 20 mg/kg calcein, respectively. 2.14. Sample collection after in vivo experiments. All the rats were sacrificed after implantation for 8 weeks. The tibias with the implants were harvested and fixed in 4% paraformaldehyde (4% PFA) for micro-CT assay and histological analysis. 2.15. Micro-CT assay. The new bone formation was determined by micro-CT (SkyScan 1176, Bruker, Madison, USA) for the fixed samples. The scanning parameters were set at 60 kV using the Al 1 mm filter with a resolution of 18 µm. After scanning, the two-dimensional (2D) and three-dimensional (3D) models were reconstructed using the NRecon software (Skyscan) and CTvol program (SkyScan) and the bone volume around the implant was determined using the DataViewer software (SkyScan) and CTAn program (SkyScan). 2.16. Histological evaluation. After micro-CT, 50 µm undecalcified sections were prepared using Exakt system (model 310 CP band system, Exakt, Oklahoma City, OK, USA) for fluorescence labeling observation under the confocal laser scanning microscope (CLSM) (Leica TCS SP8, Hamburg, Germany). The excitation/emission wavelengths used to observe chelating fluorochromes were 405/575 nm, 543/620nm and 488/520 nm for tetracycline hydrochloride (yellow), alizarin red S (red) and calcein (green), respectively. After fluorescence microscopy, the same sections were counter-stained Giemsa (MERK, Darmstadt, Germany) and examined to visualize the 12


Page 12 of 37

Page 13 of 37

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


mineralized bone tissue (in pink color). The images were captured by a fluorescence microscope (Olympus, BX53, Tokyo, Japan). For paraffin section, samples were decalcified in 10% EDTA for 4 weeks after fixation, embedded in paraffin and sectioned into 5 µm sections (Leica RM2235, Hamburg, Germany). Haematoxylin (Sigma-Aldrich, St. Louis, USA) and eosin (C0105, Beyotime Biotechonogy, Shanghai, China) (H&E) staining and Toluidine blue staining were performed to detect the specific tissue response to the implanted materials. Images were taken using a fluorescence microscope (Olympus, BX53, Tokyo, Japan). 2.17. Statistical analysis. Both in vitro and in vivo experiments were analyzed by the one-way ANOVA with Tukey’ post hoc test and expressed as means ± standard deviations (SD). A p value < 0.05 was considered to be statistically significant. 3. RESULTS AND DISCUSSION 3.1.

Characterization of PNAGA-Clay hydrogel and scaffolds. In our

experiment, Laponite XLG nanoclay was uniformly dispersed in NAGA aqueous solution. We noted that 10%, 20% and 30% NAGA concentrations were chosen to prepare hydrogels. To screen out the appropriate content for 3D printing, several contents of clay were tried. It was found that less than 7% clay content (relative to water) did not lead to a noticeable increase in the viscosity of NAGA solution, and the printed scaffold could not maintain a fixed shape with fused microstructure (Figure S1). Whereas only sticky hydrogel was formed with addition of more than 7% clay (Figure S1), possibly due to the adverse effect on photointiated polymerization. 13



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

Therefore, 7% clay was selected to add into the NAGA aqueous solution, followed by addition of photoinitiator. UV light irradiated polymerization led to the gelling of the solution. FTIR spectra evidenced the formation of polymer-clay hydrogel (Figure S2). The representative PNAGA20%-Clay hydrogel with 25.9% clay (relative to total mass of NAGA and clay) was subjected to thermal gravity analysis (TGA) (Figure S3). The results indicate that approximately 20.07% nanoclay was incorporated in the PNAGA hydrogel, suggesting that most of clay was incorporated into the gel matrix. TEM images provide direct images and local information on exfoliation and homogeneous dispersion of clay. In Figure S4, the dark platelets represent the primary particles and nanolayers of clay, and no obvious agglomeration of clay is observed. The better dispersion of clay contributes to not only high optical transparency of the matrix, but also the mechanical properties of the hydrogels, which will be discussed later.

14


Page 14 of 37

Page 15 of 37

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


Figure 1. PNAGA-Clay hydrogels exhibited extraordinary mechanical properties. Photographs of PNAGA 20%-Clay hydrogel showing its ability to withstand large deformation by compression (A and B) and elongation (C and D). Tensile stress-strain curves of PNAGA hydrogels and PNAGA-Clay hydrogels (E), and compression stress-strain curves of PNAGA hydrogels, PNAGA-Clay hydrogels and PNAGA-Clay scaffolds (PNAGA-Clay-S) (F).

The pristine PNAGA and PNAGA-Clay hydrogels were soaked in water until equilibrium swelling, and equilibrium water contents (EWCs) were determined. The EWCs of pristine and hybrid hydrogels are in the range of 72%-83.5% and 68.5-82.5%, respectively (Figure S5). The lower EWCs of PNAGA-Clay are resulted 15



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

from more stable and compact network formed by dual physical crosslinking of H-bonding and nanoclay-polymer chain linkage in the absence of conventional organic crosslinker. The mechanical properties of the fully hydrated hybrid hydrogels were measured. As presented in Figure 1, the tensile strength, Young’s modulus and elongation at break of pristine PNAGA hydrogels increase considerably with the increment of NAGA content. The corresponding maxima reach 0.76 MPa, 0.14 MPa and 800%, respectively. Notably, adding clay further contributes to the enhancement of the mechanical properties of PNAGA hydrogel. Taking PNAGA30%-Clay hydrogel as an example, its tensile strength, Young’s modulus and break stain increase up to 1.17 MPa, 0.18 MPa and 1300%, respectively. Analogously, raising NAGA content and nanoclay hybridization also lead to an evident increase of the compressive strengths of the hydrogels. Compared to the compressive strengths of pristine PNAGA10% (0.24 MPa), PNAGA20% (1.17 MPa) and PNAGA30% (2.66 MPa), the corresponding PANGA-Clay composite hydrogels achieve 0.40 MPa, 1.96 MPa and 3.50 MPa compressive strengths, respectively. Figure 1 vividly depicts that the hybrid hydrogel can withstand large strain of stretching and compression without occurrence of damage (Figure 1A-1D). The detailed data are listed out in Table S1. The marked increase in mechanical strengths of composite hydrogels results from combined dual amide hydrogen bonding interaction and physical crosslinking of nanoclay-polymer chain,28,42 as sketched in Figure. 2A.

16


Page 16 of 37

Page 17 of 37

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


Figure 2. Procedure of 3D-printing PNAGA-Clay scaffold (A) and photographs of PNAGA20%-Clay scaffolds showing its ability to resist finger compression (B and C), car wheel pressing (D, left and right denote before and after pressing) and hand folding (E and F). The scaffold is very stable even after immersing in water for a long time (G, left and right denote before and after water immersing for 3 months). The scaffolds used for car wheel pressing experiment were stained with gentian violet.

Prior to printing scaffold, we measured the viscosity of PNAGA-Clay hydrid hydrogels, one critical factor determining the printability of this ink. The viscosity of NAGA-Clay aqueous solution was about 140 Pa⋅s under 5 1/s shear rate set for 3D printing when clay concentration was 7% (different NAGA concentration did not influence the viscosity of NAGA-Clay aqueous solution a lot) (Figure S6). This 17



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

viscous fluid could be readily extruded out of the nozzle at room temperature due to shear-thinning, but quickly became highly thick upon dropping down on the printing substrate, accordingly resulting in formation of an un-collapsed structure. The printed pre-scaffold was immediately illuminated with UV light to initiate polymerization. Eventually, a high strength bioscaffold was manufactured (Figure 2A illustrates the 3D-printing procedure), and these PNAGA-Clay bioscaffolds remained excellent swelling stability in water for a long time. Figure 2 shows that even immersed in water for three months, no further swelling was noticed, suggesting that adding nanoclay into the PNAGA hydrogels did not influence hydrogen bonding interaction; even immersed in water for a much longer time, the hydrogel remained unchanged in swelling degree. This is also very important for the printed hydrogel scaffolds to maintain stable original architecture. We did not see any delamination of 10 layers of 3D scaffold immersed in water either, manifesting that multiple hydrogen bonding from dual amide in NAGA considerably enhanced the interlayer binding. It is noted that only the samples for compression could be printed with our device, so only the compression strength of the printed PNAGA-clay hydrogels was evaluated. As presented in Figure 1, the printed constructs almost remain the strengths of the original hydrogels. The fully swollen scaffold can withstand hand folding (Figure 2). And dye stained 3D scaffold was subjected to the wheel compression of a car. Amazingly, the scaffold completely recovers both in macroscopic shape and microstructure after the car rolls over it (Figure 2 and movie S1). This high strength together with swelling stability of the printed scaffold ensures its reliability both in 18


Page 18 of 37

Page 19 of 37

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


bearing external load and retaining the integrity of microstructure, which is critical for bone regeneration in vivo. 3.2.

In vitro evaluation of PNAGA-Clay hydrogel scaffolds. Next, we focused

on exploring the real application of the printed high strength hybrid hydrogels as a bioscaffold to treat animal’s bone defect (Figure 2A). Firstly, in order to inspect the neo-osteogenesis, primary rat osteoblasts (ROBs) cells were seeded onto the 3D scaffolds. CCK-8 analysis was used to test the proliferation ability of ROBs cells on the scaffolds. After one-day culturing, no obvious difference is observed among three different scaffolds. We can see the cells adhere well and spread out on all the scaffolds (Figure 3A). After 7 days, both PNAGA10%-Clay and PNAGA20%-Clay hydrogel scaffolds exhibit good cytocompatibility (Figure 3A). This is further confirmed by Live-dead assay in Figure 3B(III) where almost no dead cells can be found. Nevertheless, for PNAGA30%-Clay hydrogel, the growth rate of ROBs cells decreases and slight toxicity is detected. We cannot give a clear explanation. A probable reason is that the excessive crosslinking density is unfavorable for the growth of the cells. SEM and CLSM images of PNAGA20%-Clay scaffolds with cell culture for 7 days clearly exhibit that ROBs cells almost cover the whole holes and penetrate into the inside of the 3D scaffold, and the cells spread very well in spindle shape (Figure 3B). As a biocompatible material, nanoclay contains several kinds of ions, which can diffuse out to influence the cell behavior.43 Here, we mainly traced the release of silicon ions and magnesium ions, two bioactive ions that have been demonstrated to elicit osteogenesis.44-45 ICP results present that these two ions diffuse 19



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

out of the hydrogels gradually, and 58-85 µg/mL Mg2+ and 110-150 µg/mL Si4+ that are both in the effective concentrations for eliciting bone regeneration can be detected at 21 days for the composite hydrogels (Figure 4).44-45 The release rate increases steadily for all samples at the first 7 days, and then the release slows down. A slower release rate is observed for composite hydrogels with higher NAGA content. This means that higher crosslinking density hindered the diffusion of ions. As shown in Figure S7, the average pore sizes of the PNAGA10%-Clay, PNAGA20%-Clay and PNAGA30%-Clay are 0.96 µm, 0.48 µm and 0.14 µm, respectively. Clearly, as the concentration of NAGA increases, the pore size for the microstructure of the PNAGA-Clay hydrogels decreases significantly due to the higher crosslinking density. Since the major application of our scaffolds is bone tissue engineering, what we are concerned with here is whether the key ions release from the scaffolds can promote osteogenic differentiation, and thus accelerate new bone formation.

20


Page 20 of 37

Page 21 of 37

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


Figure 3. In vitro evaluation of PNAGA-Clay scaffolds. CCK-8 analysis of ROBs cells after cultured on PNAGA-Clay scaffolds for 1, 3, 5 and 7 days (A). SEM images of ROBs cells growing within PNAGA 20%-Clay scaffold (B(I) and B(II), cells are colored with pink). Cells grow into the hollow of the scaffold (B(II)). Live-dead assay and cytoskeleton staining of ROBs cells after seeded onto PNAGA 20%-Clays scaffold for 7 days (B(III) and B(IV)). Osteogenic gene expression (ALP, OCN and collagen) of ROBs cells cultured on PNAGA 20% hydrogel and PNAGA 20%-Clay hydrogel at day 7 (C). ALP staining of ROBs cells after seeded within PNAGA-Clay for 21 days. D(I): PNAGA 10%-Clay, D(II): PNAGA 20%-Clay, D(III): PNAGA 30%-Clay. ALP analysis of ROBs cells after cultured within PNAGA-Clay scaffolds for 4, 7, 14 and 21 days (D(IV)). Asterisks (*) denote significant differences (*p

3D-Printed High Strength Bioactive ... - ACS Publications

Recommend Documents