chitosan hydrogels for the treatment of vertebral compression fractures

International Journal of Biological Macromolecules 130 (2019) 88–98

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International Journal of Biological Macromolecules journal homepage: http://www.elsevier.com/locate/ijbiomac

Inductive co-crosslinking of cellulose nanocrystal/chitosan hydrogels for the treatment of vertebral compression fractures Soheila Ali Akbari Ghavimi a, Ethan S. Lungren a, Trent J. Faulkner a, Mary A. Josselet a, Ying Wu a, Yisheng Sun a, Ferris M. Pfeiffer a, Christina L. Goldstein b, Caixia Wan a, Bret D. Ulery a,⁎ a b

Department of Biomedical, Biological, and Chemical Engineering, University of Missouri, Columbia, MO 65211, United States of America Department of Orthopedic Surgery, University of Missouri, Columbia, MO, United States of America

a r t i c l e

i n f o

Article history: Received 16 November 2018 Received in revised form 11 February 2019 Accepted 14 February 2019 Available online 16 February 2019 Keywords: Vertebral compression fractures Hydrogel Chitosan Cellulose nanocrystals

a b s t r a c t Vertebral compression fractures are a very common consequence of osteoporosis for which injection of a nonbiodegradable, non-bioactive, mechanically-stiff polymer bone cement into the vertebral body is the most common treatment. Recently, there has been growing interest in using bioactive, degradable, and bone biomechanics-matching products as an alternative approach for treating these fractures. In this research, we focused on creating injectable, chitosan-based hydrogels that can convey mechanical strength similar to vertebral bone as well as possess inherent osteoinductivity. First, we investigated the effects of three different factors – 1) bioactive phosphate ionic crosslinking; 2) genipin covalent crosslinking; 3) mechanically reinforcing cellulose nanocrystal incorporation – on the material properties of chitosan-based hydrogels. Mesenchymal stem cells were then exposed to hydrogels with optimum mechanical properties and stability in order to assess the biological effects of the bioactive phosphate ionic crosslinker. Our results show that hydrogels with higher ionic and covalent crosslinking ratios supplemented with neutral cellulose nanocrystals possessed desirable compressive strength and stability. Also, the significant osteoinductivity of these composite hydrogels demonstrated their potential to function as an injectable system for the future treatment of vertebral compression fractures. © 2019 Elsevier B.V. All rights reserved.

1. Introduction Vertebral compression fractures (VCFs) are caused by the collapse of vertebral bone into itself, commonly due to trauma [1], and is one of the leading cause of disability in the elderly [2–4]. Approximately 1.5 million VCFs occur annually in the United States, primarily as a consequence of osteoporosis [5]. These fractures usually occur in the mid-thoracic or thoracolumbar transition zones of the spine and cause sudden focal back pain and loss of function [6,7]. In complicated cases with pressure being exerted on the spinal cord, VCFs can cause compromised neurological function including weakness, loss of sensation in the lower extremities, and lack of control of the bladder and bowel [8]. In addition, VCFs can cause pulmonary dysfunction, loss of mobility, chronic spinal deformity, chronic pain, and depression [9]. Vertebroplasty and kyphoplasty are minimally invasive treatments for VCFs [1,10–12] which involve modifying fractured bone tissue through the injection of poly(methylmethacrylate) (PMMA) bone cement into the vertebral body [13]. Short-term success has been achieved with this approach, however, there are considerable concerns regarding the use of PMMA. Due to the exothermic nature of its ⁎ Corresponding author. E-mail address: [email protected] (B.D. Ulery).

https://doi.org/10.1016/j.ijbiomac.2019.02.086 0141-8130/© 2019 Elsevier B.V. All rights reserved.

polymerization, PMMA can cause post-injection thermal damage to the surrounding tissue. Additionally, PMMA is non-bioactive and nonbiodegradable, so it cannot facilitate regeneration of the damaged bone nor be removed naturally by the body over time. PMMA is also about twice as dense and possesses a much higher compressive strength than native bone tissue which causes a significant risk of new adjacent fractures in patients treated with this material [14]. Due to these limitations, new materials for the treatment of VCFs are greatly needed. Chitosan-based hydrogels have been well studied as an injectable, biodegradable, and bioactive system for bone regenerative applications [15,16]. The cationic nature of chitosan due to the presence of pendant amine groups in its molecular structure makes it water soluble under mildly acidic conditions [17,18]. This feature also allows for chitosan to be able to be in situ crosslinkable into a hydrogel, making it a suitable candidate for non-invasive or minimally-invasive delivery. Chitosan hydrogels are elastic and rubbery, similar to natural tissue and easily adopt the geometry of the void space they are supplied. One disadvantage of chitosan hydrogels are their poor mechanical properties though this can be overcome by compositing them with other materials [19,20]. Recently, nanomaterials have been widely adopted as hydrogel reinforcing agents [21–23]. Cellulose nanocrystals (CNCs) in particular have received a lot of attention as a reinforcing agent due to their unique

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properties such as large surface area, high aspect ratio, adequate mechanical strength, low density, biocompatibility, and biodegradability [24–27]. Physically incorporating CNCs into chitosan hydrogels has the potential to enhance the mechanical properties of the resulting composite for VCF treatment [28]. The terminal hydroxyl groups on CNCs provide the opportunity for their easy chemical modification, allowing for the generation of CNCs with different surface charges. Chitosan hydrogels can be fabricated through physical or chemical crosslinking or a combination of the two. Due to the cationic amine groups of chitosan, ionic crosslinking can be achieved between chitosan and small anionic molecules, such as sulfates, carbonates, and phosphates [29–31]. Ionic interactions can lead to different properties depending on the charge density and size of the ionic molecule [29]. As ionic crosslinking does not require the presence of a potentially toxic covalent crosslinkers, it is considered safe for clinical application [32]. Further, phosphate-based ionic crosslinkers has the advantage of potentially being osteoinductive since phosphate is known to function as a signaling molecule capable of facilitating the differentiation of stem cells into osteoblasts [33,34]. However, the weak mechanical strength and the uncontrolled dissolution of ionically crosslinked hydrogels limits their potential for load bearing applications [32]. Covalent crosslinking produces irreversible permanent hydrogels by creating stiff physical bonds using amine, hydroxyl, and carboxylate groups. Many bifunctional small molecules have been used to covalently crosslink chitosan including glutaraldehyde, diglycidyl ether, diisocyanate, and genipin [19,35–39]. In general, covalent crosslinking can greatly improve mechanical properties compared to ionic crosslinking. Unfortunately, a major drawback of most of these crosslinkers is that they have been found to possess moderate to severe cytotoxicity [40]. Genipin, a naturally derived molecule from plants in the Gardenia genus, has been found to be one of the only biocompatible covalent crosslinkers currently available [41]. In specific, it has been shown to be an effective crosslinking agent for amine containing polymers with significantly less cytotoxicity than the commonly utilized glutaraldehyde [42,43]. Although hydrogels are hydrophilic, their rapid dissolution is prevented through crosslinking and/or physical entanglement [43]. The physical properties of hydrogels are highly regulated by the crosslinking type (i.e. ionic or covalent) and density [44,45]. Improved mechanical properties and chemical stability can be readily achieved with covalent crosslinking [43,45,46] whereas high swelling ratio and degradability can be easily conveyed with ionic crosslinking [45–47]. Also, increased crosslinker density has been found to be correlated with hydrogel mechanical strength and stability [40,48]. We believe that chitosan hydrogels crosslinked both ionically (to increase hydrogel swelling) and covalently (to enhance mechanical properties and stability) as well as reinforced with CNCs (to improve hydrogel mechanical properties) can serve as a multi-component composite biomaterial that can overcome the limitations associated with PMMA cements for the treatment of VCFs. The content of each component and their physical and chemical interactions can tune hydrogel mechanical properties and stability as well as osteoinductivity. In this report, the effect of three different factors – the ratio and type of ionic crosslinker, the ratio of covalent crosslinker, and the incorporation of CNCs with different surface charge – on the mechanical properties and stability of chitosan hydrogels were investigated. Hydrogels with better mechanical properties and stability were then exposed to mesenchymal stem cells to evaluate the osteoinductivity of these materials. 2. Experimental

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CNCs (−CNCs), hydrochloric acid/citric acid hydrolysis of CMCs with no further modification was employed [49]. In specific, 3 g of CMC was added into 150 mL of the mixture of citric and hydrochloric acids that was preparing by mixing 3 M citric acid and 6 M hydrochloric acid with a volumetric ratio of 9:1. The hydrolysis reaction was first conducted at 80 °C with continuous stirring for 4 h. Thereafter, the reaction mixture was transferred into dialysis tubes for washing in DI water. The asprepared −CNCs were also utilized to synthesize neutral CNCs (0CNC) by capping −CNCs with cetyltrimethyl ammonium bromide (CTAB) [50]. First, the −CNCs concentration was adjusted to 0.1 wt% and its pH adjusted to 10 with NaOH solution. Next, CTAB powder was added to the CNC suspension by using a molar ratio of 1:2 of the carboxyl groups of CNCs to CTAB. The reaction was then initiated at 60 °C for 3 h with continuous stirring and continued overnight at room temperature. Dialysis was also applied to wash the samples until neutral pH was achieved. Positively-charged CNCs (+CNCs) were synthesized by sulfuric acid hydrolysis of CMCs followed by capping with epoxy propyltrimethyl allyl chloride (EPTMAC) [51]. The CNCs were first obtained by sulfuric acid hydrolysis. Briefly, 1 g of CMC was hydrolyzed by 10 mL of 64% sulfuric acid under 45 °C for 30 min. The CNCs were first dialyzed within DI water before reacting with EPTMAC. EPTMAC modification was catalyzed by 0.5 M NaOH (1 g NaOH/10 g CNC) and the amount of EPTMAC was determined by a molar ratio of 3:1 or EPTMAC over the anhydroglucose units of CNCs. The reaction lasted for 6 h under 65 °C with mild stirring. At the end of the reaction, 95% ethanol was added to the mixture by a 1:10 volumetric ratio to quench the unreacted chemicals. The +CNCs were then washed by DI water within dialysis tubes. 2.2. Hydrogel preparation Low molecular weight chitosan (MW ~ 50,000–190,000 Da - Sigma Aldrich) was dissolved in 0.5% acetic acid in double-distilled water (ddH2O) solution at 1.4% wt/v using constant magnetic stirring for 72 h at room temperature and vacuum filtered after dissolution. CNCs with three different surface charges (−CNC, 0CNC, or +CNC) were suspended in chitosan solution at 0.07% wt/v (5% wtCNC/wtChitosan) by employing 10 min of sonication using a probe-tip sonicator (Thermo Fisher) at 200 W. Chitosan solution without any CNCs (noCNC) was utilized as a control group. To prepare crosslinking solutions, ionic crosslinkers (i.e. disodium phosphate and sodium bicarbonate) and covalent crosslinker (i.e. genipin) were dissolved in ddH2O using the probe-tip sonicator at 200 W. The crosslinking solutions were created so that for each hydrogel the molar ratio of crosslinker per deacetylated sites of chitosan, that are available for crosslinking (85%), was 5, 10, and 15 for ionic crosslinkers and 0.312, 1.25, and 2.5 for covalent crosslinker. Hydrogels were made by adding the appropriate crosslinking solution to a CNC/chitosan solution and mixing using high-speed vortexing for 15 s. The samples were then placed in a water bath at 37 °C for further evaluation. Crosslinking solution combinations resulted in hydrogels with ratios of 5:0.312:1, 5:1.25:1, 5:2.5:1, 10:0.312:1, 10:1.25:1, 10:2.5:1, 15:0.312:1, 15:1.25:1, or 15:2.5:1 of phosphate:genipin:chitosan (P:G:C) or carbonate:genipin:chitosan (C:G:C). Different hydrogel samples, their components, and labels are summarized in Table 1. 2.3. Fourier transform infrared (FTIR) spectroscopy FTIR spectra were recorded on a Nicolet 6700 FT-IR (Thermo Scientific). Spectra were obtained in the range of 4000 cm−1 to 400 cm−1 with the purpose of evaluating chemical structure changes, especially the ratio of the nitrogen‑hydrogen bonds (N\\H) to carbon‑nitrogen (C\\N) bonds after crosslinking.

2.1. Synthesis of cellulose nanocrystals 2.4. Gelation time measurements Cellulose nanocrystals (CNCs) with three different surface charge, namely negative, neutral, or positive, were obtained from cellulose microcrystals (CMCs, Avicel PH-101). To prepare negatively-charged

Gelation time was determined using the test tube inversion method [31]. Briefly, the hydrogel mixture was incubated in a vial at 37 °C which

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Table 1 Hydrogel component and labels. All the contents are reported for each representative hydrogel product. *CNC can include noCNC, −CNC, 0CNC, or +CNC for *CNC/P:G:C hydrogels and can possess noCNC and 0CNC for *CNC/C:G:C hydrogels. For no genipin and high concentrations of genipin (i.e. 0:1, 5:1, and 7.5:1) only 0CNC were utilized. The unit mol:mol for crosslinkers (i.e. phosphate, carbonate, and genipin) is moles of the crosslinker to moles of chitosan deacetylated site that are available for crosslinking. Water Chitosan *CNC Phosphate Carbonate Genipin Label (ml) (mg) (mg) mol:mol mol:mol mol:mol 1.2 1.2 1.2 1.2

14 14 14 14

0.7 0.7 0.7 0.7

5:1 5:1 5:1 10:1

0:1 0:1 0:1 0:1

0.312:1 1.25:1 2.5:1 0.312:1

1.2 1.2 1.2 1.2

14 14 14 14

0.7 0.7 0.7 0.7

10:1 10:1 15:1 15:1

0:1 0:1 0:1 0:1

1.25:1 2.5:1 0:1 0.312:1

1.2 1.2 1.2 1.2 1.2

14 14 14 14 14

0.7 0.7 0.7 0.7 0.7

15:1 15:1 15:1 15:1 0:1

0:1 0:1 0:1 0:1 5:1

1.25:1 2.5:1 5:1 7.5:1 0.312:1

1.2 1.2 1.2

14 14 14

0.7 0.7 0.7

0:1 0:1 0:1

5:1 5:1 10:1

1.25:1 2.5:1 0.312:1

1.2

14

0.7

0:1

10:1

1.25:1

1.2 1.2

14 14

0.7 0.7

0:1 0:1

10:1 15:1

2.5:1 0.312:1

1.2

14

0.7

0:1

15:1

1.25:1

1.2

14

0.7

0:1

15:1

2.5:1

*CNC/P:G:C 5:0.312:1 *CNC/P:G:C 5:1.25:1 *CNC/P:G:C 5:2.5:1 *CNC/P:G:C 10:0.312:1 *CNC/P:G:C 10:1.25:1 *CNC/P:G:C10:2.5:1 *CNC/P:G:C 15:0:1 *CNC/P:G:C 15:0.312:1 *CNC/P:G:C 15:1.25:1 *CNC/P:G:C 15:2.5:1 *CNC/P:G:C 15:5:1 *CNC/P:G:C 15:7.5:1 *CNC/C:G:C 5:0.312:1 *CNC/C:G:C 5:1.25:1 *CNC/C:G:C 5:2.5:1 *CNC/C:G:C 10:0.312:1 *CNC/C:G:C 10:1.25:1 *CNC/C:G:C10:2.5:1 *CNC/C:G:C 15:0.312:1 *CNC/C:G:C 15:1.25:1 *CNC/C:G:C 15:2.5:1

was then inverted every 30 s. The gelation time was recorded as the time that the mixture no longer flowed anymore and the vial could be completed inverted without the sample falling due to gravity. 2.5. Compressive strength measurement Crosslinked cylindrical hydrogel samples with thicknesses of ~5 mm and diameters of ~10 mm were removed from their gelation vials. The mechanical properties of the samples were measured via compression test using a TA Universal mechanical test device. The strain response for each sample was monitored under a 5 kg load and a crosshead speed of 1 mm/min until 80% strain is achieved. 2.6. Swelling measurements Hydrogel swelling was assessed by first immersing samples in phosphate buffered saline (PBS) at 37 °C for 24 h. After 24 h, the samples were removed from PBS and blotted gently with Kimwipes to remove surface associated water and weighed. The samples were then frozen at −80 °C and lyophilized under vacuum (0.1 mm Hg) and at −50 °C for 72 h to make sure all the solvent was removed. The samples were weighed after 3 days and the swelling ratio was measured using Eq. (1):

Swelling ¼

Ww Wd

ð1Þ

where Ww is the wet hydrogel weight before lyophilization and Wd is the dry hydrogel weight after lyophilization.

2.7. Mass loss measurement Mass loss measurements were conducted by immersing hydrogel samples in PBS at 37 °C for a period of 14 days. At specific intervals, the samples were removed from PBS and blotted gently with Kim wipes to remove surface associated water and weighed. Mass loss (%) was determined using the Eq. (2): Mass loss ð%Þ ¼

W 0 −W t W0

ð2Þ

where W0 and Wt are initial weight and weight at the specific time point, respectively. The endpoint data (i.e. day 14) are reported in this manuscript. 2.8. Cell culture and seeding Mesenchymal stem cells (MSCs) were purchased from Cyagen and initially cultured in T-75 cell culture flasks (Corning) and grow in Dulbeco's modified Eagle's medium (DMEM, Invitrogen), supplemented with 10% fetal bovine serum (FBS, Invitrogen) and 1% Penicillinstreptomycin (Pen-Strep, Invitrogen) at 37 °C in a humidified incubator with 5% CO2. Media was changed every 48 h until cells approached ~80% confluency after which they were dissociated using a 0.05% trypsinEDTA solution (Invitrogen). Detached MSCs were counted with a hemocytometer and passed to new T-75 flasks at a splitting ratio of 1:4 or 1:5 dependent on overall cell count. After the 5th passage, cells were used for in vitro bioactivity studies. Tissue cultured polystyrene 24-well plates (Corning) were seeded with 30,000 cells/well and exposed to growth media alone as a negative control, 0CNC/P:G:C hydrogels with different ionic crosslinking ratios (i.e. 5:2.5:1, 10:2.5:1, 15:2.5:1) as potentially bioactive hydrogels, and 0CNC/C:G:C hydrogels with different ionic crosslinking ratios (i.e. 5:2.5:1, 10:2.5:1, 15:2.5:1) as potentially inert hydrogels. MSCs were cultured for up to 14 days at 37 °C in a humidified incubator with 5% CO2 with growth media changed every 3 days. Cell proliferation, viability, alkaline phosphatase (ALP) activity, and mineralization were assessed at 1, 3, 7, and 14 days. 2.9. Proliferation assay Cell proliferation was determined using the Quanti-iT PicoGreen dsDNA Assay (Thermo Fisher Scientific). At each endpoint, the samples were rinsed with PBS and exposed to 1% Triton X-100 (Sigma-Aldrich) followed by three freeze-thaw cycles in order to lyse the cells. Lysates were diluted with TE buffer (200 mM Tris-HCL, 20 mM EDTA, pH 7.5) and mixed with PicoGreen reagent according to the manufacturer's protocol. A BioTek Cytation 5 fluorospectrometer plate reader was utilized to measure the fluorescence of each sample (ex. 480 nm, em. 520 nm) and the cell number was determined using a MSC standard curve (0–200,000 cells/mL). 2.10. Viability assay Cell viability was evaluated at each time point using a MTS Cell Proliferation Colorimetric Assay Kit (BioVision). MTS reagent (20 μL) was added to growth media (500 μL) followed by 4 h incubation at 37 °C in a humidified incubator with 5% CO2. Absorbance of each sample was measured at 490 nm. Cell viability was reported as the ratio of absorbance in the experimental groups compared to exposure to only growth media as a negative control. 2.11. Alkaline phosphatase activity assay Cell ALP activity was quantified at each time point using an Alkaline Phosphatase Assay Kit (BioVision). In brief, 20 μL of cell lysate was combined with 50 μL of p-nitrophenyl phosphate (pNPP) solution in assay

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buffer. The mixture was incubated for 1 h at room temperature away from light. The reaction was stopped by the addition of 20 μL of stop solution and the absorbance of the resulting solution was measured at 405 nm using a plate reader. To eliminate any background effects, 1% Triton X-100 was incubated with pNPP, exposed to stop solution after 1 h, and its absorbance deducted from the value for each sample. This was converted to the content of p-nitrophenyl (pNP) using a standard curve (0–20 nmol/mL) which was dephosphorylated using excess ALP enzyme. ALP activity was reported as pNP content normalized by cell count. 2.12. Mineralization assay Cell-based mineral deposition was measured at each time point using an Alizarin Red assay. At each time point, the media was removed after which the cells were washed with ddH2O and fixed in 70% ethanol for 24 h. The ethanol was removed and the samples were incubated in 1 mL of 40 mM Alizarin Red solution (Sigma-Aldrich) for 10 min. The samples were rinsed with ddH2O several times to make sure all nonabsorbed stain was removed. Absorbed Alizarin Red was desorbed using 1 mL of a 10% cetylpyridinium chloride (CPC, Sigma-Aldrich) solution after which stain concentration was measured at 550 nm using a plate reader. Absorbance of each sample was converted to the concentration of absorbed Alizarin Red using a standard curve (0–0.2740 mg/mL). Samples above the standard curve linear range were diluted with CPC until a solution in the linear range was obtained. Cell-based mineral deposition was calculated by subtracting mineralization found in blank wells exposed to the same experimental conditions. All results were normalized by cell count. 2.13. Statistical analysis JMP software was used to make comparisons between experimental groups with Tukey's HSD test specifically utilized to determine pairwise statistical differences (p b 0.05). The statistical analysis results are reported in the supporting information section. Groups that possess different letters have statistically significant differences in mean whereas those that possess the same letter have means that are statistically insignificant in their differences. 3. Results FTIR analysis of the hydrogels was used to confirm covalent crosslinking via the increase in the peak deviation intensity ratio of

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C\\N bonds to N\\H bonds found in different samples. The characteristic peaks of chitosan are N\\H/O\\H, C\\H, N\\H, C\\N, and C\\O\\C which absorb at approximately 3400, 2800, 1658, 1425, and 1076 cm−1, respectively. Since genipin crosslinking occurs through covalent bonding N\\H groups, an increase in the intensity ratio of C\\N/N\\H is expected after gelation. As shown in Fig. S1, the baseline C\\N/N\\H ratio for chitosan alone is 0.77 as compared to the increased 1.18 observed for 0CNC/P:G:C 15:2.5:1. Increased covalent crosslinker content led to increased C-N/N-H ratio ranges (i.e. 0.94–0.95 for X:0.312:1, 1.05–1.07 for X:1.25:1, and 1.18 for X:2.5:1). In contrast, altering ionic crosslinker type or ratio as well as CNC inclusion and surface charge had no noticeable impact on C\\N/N\\H ratios indicating these components did not participate in nor impact covalent crosslinking. 3.1. Materials properties of cellulose nanocrystal/phosphate:genipin:chitosan (CNC/P:G:C) hydrogels Gelation times for different CNC/P:G:C hydrogels (Fig. 1) show that this material property is directly influenced by ionic crosslinking ratio, covalent crosslinking ratio, CNC presence, and CNC surface charge. In specific, increasing the phosphate:chitosan ratio generally decreased gelation time. The same trend was observed for the genipin:chitosan ratio, however, the effect of the covalent crosslinking ratio was more pronounced than that of the ionic crosslinking ratio. CNC surface charge also influenced gelation time with the shortest to longest times typically being +CNC b noCNC b −CNC ≤ 0CNC regardless of ionic and covalent crosslinking ratio. For all high genipin to chitosan ratio (2.5:1) hydrogels, gelation occurred in b1 min. Gelation times for all samples except 0CNC/P:G:C 5:0.312:1 were found to be b30 min. The swelling ratio and mass loss of CNC/P:G:C hydrogels are reported in Table 2. Hydrogel swelling ratios varied slightly as a function of hydrogel components. Increasing ionic and/or covalent crosslinking ratio caused the swelling ratio for all hydrogels to decrease regardless of CNC surface charge. The presence of hydrogels with CNC incorporation possessed lower swelling ratios compared to hydrogels without CNCs, however, the surface charge of CNCs did not considerably influence the swelling ratio. Ionic and covalent crosslinking ratios were found to not remarkably influence hydrogel dissociation after 14 days of incubation. In contrast, CNC surface charge can significantly impact mass loss over this time period. Hydrogels without any CNCs had the greatest mass loss with all samples losing N80% of their mass. Addition of −CNC decreased mass loss to about 70% for all P:G:C ratios. The mass loss for hydrogels containing +CNC dropped further to 50 to 70% regardless of formulation. 0CNC had the most significant influence on

Fig. 1. Gelation time for P:G:C hydrogels with different crosslinking ratios, CNC incorporation, and CNC surface charge. P:G:C hydrogels were prepared without CNCs (X) or with CNCs with negative surface charge (−), no surface charge (O), or positive surface charge (+). The inversion method at 37 °C was used to determine in situ gelation time. Results are an average of four independent measurements.

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Table 2 Swelling ratio and mass loss for P:G:C hydrogels with different crosslinking ratios, CNC incorporation, and CNC surface charge. Mass loss was measured by immersing hydrogels in PBS at 37 °C for 14 days and comparing the change in mass over time. The swelling ratio was determined as the ratio of wet weight to dry weight of hydrogels after immersion in PBS for 24 h at 37 °C. Results are an average of four independent measurements. P:G:C

5:0.312:1 5:1.25:1 5:2.5:1 10:0.312:1 10:1.25:1 10:2.5:1 15:0.312:1 15:1.25:1 15:2.5:1

No

CNC

0

CNC

Swelling ratio

Mass loss (%)

Swelling ratio

32.3 ± 1.5 30.0 ± 5.4 29.3 ± 4.4 24.4 ± 3.7 23.4 ± 6.0 18.2 ± 5.0 19.9 ± 2.2 19.0 ± 1.4 18.2 ± 3.1

93.7 ± 2.1 88.2 ± 6.2 88.0 ± 6.2 90.5 ± 2.4 99.9 ± 0.8 93.0 ± 3.2 90.9 ± 1.6 89.7 ± 4.5 83.5 ± 9.8

28.1 ± 9.0 22.8 ± 4.0 20.3 ± 5.1 20.3 ± 0.9 18.7 ± 1.4 16.3 ± 3.4 17.4 ± 1.3 16.3 ± 0.9 16.6 ± 4.2

+

CNC

Mass loss (%)

Swelling ratio

68.3 ± 20.4 65.6 ± 19.81 69.7 ± 16.8 69.8 ± 12.4 70.1 ± 3.9 69.4 ± 6.6 71.2 ± 5.6 69.5 ± 11.2 69.9 ± 14.4

hydrogel stability yielding hydrogels that only lost around 25–35% of their mass except for the formulation with the lowest ionic and covalent crosslinking ratios (i.e. 0CNC/P:G:C 5:0.312:1). The compressive strength of P:G:C hydrogels are outlined in Fig. 2 for which statistical analysis was conducted to determine the influence of ionic crosslinking ratio (Table S1), covalent crosslinking ratio (Table S2), and CNC incorporation and surface charge (Table S3). Altering the phosphate to chitosan crosslinking ratio was found to have a minimal impact on compressive strength whereas increasing the genipin to chitosan crosslinking ratio significantly improved hydrogel mechanics. Also, the incorporation of 0CNCs enhanced the mechanical strength of P:G:C hydrogels whereas the inclusion of charged CNCs ( −CNCs or +CNCs) did not influence this materials property. The greatest compressive strength observed among the different formulations was 109 kPa which was found with 0CNC/P: G:C 15:2.5:1. 3.2. Effect of covalent crosslinking ratio on hydrogel materials properties The materials properties of chitosan hydrogels with a 15:1 ionic crosslinker ratio and 0CNC supplementation were found to be the best for the desired application, but could be further improved. As genipin content was found to have a significant influence on hydrogel mechanical properties, the effect of covalent crosslinking ratio were further investigated using 5:1 and 7.5:1 genipin:chitosan. The gelation time for 0 CNC/P:G:C 15:X:1 hydrogels with varying covalent crosslinking ratio

23.9 ± 0.4 19.8 ± 0.6 17.9 ± 1.2 19.4 ± 2.5 15.7 ± 0.8 14.3 ± 2.2 18.7 ± 0.8 12.9 ± 0.6 10.6 ± 1.9

CNC

Mass loss (%) 48.0 ± 4.5 32.3 ± 2.2 34.1 ± 14.9 29.8 ± 8.6 26.3 ± 5.6 32.9 ± 2.1 31.4 ± 10.1 27.7 ± 12.9 34.2 ± 14.9

Swelling ratio 26.5 ± 0.9 22.7 ± 3.0 20.9 ± 4.3 21.7 ± 4.2 17 ± 7.9 15.4 ± 4.0 18.8 ± 3.1 16.7 ± 5.4 13.6 ± 1.6

Mass loss (%) 71.2 ± 0.9 64.0 ± 11.6 59.3 ± 14.5 64.4 ± 2.1 55.3 ± 29.9 64.0 ± 5.8 61.1 ± 10.5 58.5 ± 4.2 47.5 ± 6.6

are found in Fig. 3. Similar to previous results, gelation time was greatly dependent on covalent crosslinking ratio with 0CNC/P:G:C 15:5:1 and 0 CNC/P:G:C 15:7.5:1 hydrogels solidifying in b30 s. The effect of covalent crosslinking content on the swelling ratio and mass loss of 0CNC/P:G:C 15:X:1 hydrogels is reported in Table 3. Without the presence of genipin, chitosan hydrogels undergo more limited swelling with a ratio of 8.3 and rapid erosion leading to nearly 90% mass loss in 14 days. Interestingly, high covalent crosslinking ratios (i.e. 5:1 and 7.5:1) also yield more limited swelling and greater mass loss than the previously investigated ratios (i.e. 0.312:1, 1.25:1, and 2.5:1). In fact, 0CNC/P:G:C 15:0:1 hydrogels and 0CNC/P:G:C 15:7.5:1 hydrogels possess almost the same swelling ratio and mass loss. The effect of covalent crosslinking on the compressive strength and stress-strain curves of hydrogels under compression are demonstrated in Fig. 4. A positive correlation between hydrogel compressive strength and genipin:chitosan ratio was observed (Fig. 4a). The compressive strengths of 0CNC/P:G:C 15:5:1 and 0CNC/P:G:C 15:7.5:1 hydrogels were 183 kPa and 255 kPa, respectively, both statistically significantly greater than all other formulation evaluated. Stress-strain curves revealed that genipin:chitosan ratio also impacts the deformation behavior of the hydrogels (Fig. 4b). Young's Modulus for compression was found to be positively correlated with genipin content. Although all hydrogels underwent an elastic deformation followed by a plastic deformation, 0CNC/P:G:C 15:5:1 and 0CNC/P:G:C 15:7.5:1 hydrogels showed more limited plastic deformation than all formulations with lower genipin:chitosan ratios (i.e. 0:1 to 2.5:1). Additionally, these higher

Fig. 2. Compressive strength for P:G:C hydrogels with different crosslinking ratios, CNC incorporation, and CNC surface charge. P:G:C hydrogels were prepared without CNCs (X) or with CNCs with negative surface charge (−), no surface charge (O), or positive surface charge. Results are an average of four independent measurements. Statistical analysis of the differences between groups are available in supplementary information (Table S1, Table S2, and Table S3).

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Fig. 3. Gelation time for 0CNC/P:G:C 15:X:1 hydrogels with different covalent crosslinking ratios. The inversion method at 37 °C was used to determine in situ gelation time. Results are an average of four independent measurements.

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The swelling ratio and mass loss of CNC/C:G:C hydrogels are reported in Table 4. Hydrogel swelling ratios varied slightly as a function of hydrogel components. Increasing carbonate or genipin content caused a slight decrease in hydrogel swelling. Additionally, the incorporation of 0CNCs also yielded lower hydrogel swelling ratios than those that did not have CNCs. Similar to P:G:C hydrogels (Table 2), ionic and covalent crosslinking did not dramatically alter C:G:C hydrogel mass loss over 14 days which was found to by 55% to 65% regardless of formulation. Addition of 0CNCs was much more impactful, decreasing hydrogel dissociation to only about 35% to 45% over the same time period. The compressive strength of C:G:C hydrogels are detailed in Fig. 6 for which statistical analysis was conducted to determine the impact of ionic crosslinking ratio (Table S1), covalent crosslinking ratio (Table S2), and CNC incorporation (Table S3). The compressive strength of chitosan hydrogels was not significantly influenced by changes in carbonate or genipin content nor the presence of 0CNCs generally. 0CNC/C: G:C hydrogels did possess have statistically significant improvements in their compressive strength over the noCNC/C:G:C hydrogels for C:G:C ratios of 5:0.312:1, 15:0.312:1, and 15:2.5:1. The greatest compressive strength observed among the different formulations was 53 kPa which was found with 0CNC/5:2.5:1C:G:C hydrogels.

genipin:chitosan ratio hydrogels fractured at much lower strains (i.e. around 100%) than the others hydrogels (i.e. 250% to 310%).

3.3. Materials properties of cellulose nanocrystal/carbonate:genipin:chitosan hydrogels Since phosphate is known to be an osteoinductive molecule, inert carbonate has been chosen as an alternative ionic crosslinker to serve as a bioactivity negative control. As the charge density in phosphate and carbonate are different, the impact carbonate ionic crosslinking has on chitosan hydrogels materials properties was explored. As 0CNC/ P:G:C hydrogels were found to have the best compressive strength, only 0CNC/C:G:C hydrogels were investigated. Gelation time for different 0CNC/C:G:C hydrogels are shown in Fig. 5 and this materials property was found to depend on ionic crosslinking ratio, covalent crosslinking ratio, and 0CNC incorporation. Carbonate content was found to impact gelation time for 0CNC/C:G:C hydrogels but not noCNC/C:G:C hydrogels. Increasing genipin content was found to decrease gelation time for C:G:C regardless of 0CNC inclusion or not. 0CNC incorporation retarded gelation time for some C:G:C hydrogels with the greatest impact seen with low ionic crosslinking ratio (i.e. 5:1) or high covalent crosslinking ratio (i.e. 2.5:1). These trends are quite similar to what was observed with P:C:G hydrogels (Fig. 1).

Table 3 Swelling ratio and mass loss for 0CNC/P:G:C 15:X:1 hydrogels with different covalent crosslinking ratios. Mass loss was measured by immersing hydrogels in PBS at 37 °C for 14 days and comparing the change in mass over time. The swelling was determined as the ratio of wet weight to dry weight of hydrogels after immersion in PBS for 24 h at 37 °C. Results are an average of four independent measurements. P:G:C

0

CNC

Swelling ratio 15:0:1 15:0.312:1 15:1.25:1 15:2.5:1 15:5:1 15:7.5:1

8.3 ± 0.1 18.7 ± 0.8 12.9 ± 0.5 10.6 ± 1.9 10.2 ± 3.5 9.1 ± 0.6

Mass loss (%) 89.7 ± 8.8 31.4 ± 10.1 27.7 ± 3.4 34.2 ± 4.8 54.7 ± 6.7 84.4 ± 12.1

Fig. 4. Compressive strength (a) and stress-strain curves (b) for 0CNC/P:G:C 15:X:1 hydrogels with different covalent crosslinking ratios. Results are an average of four independent measurements. Groups that possess different letters have statistically significant differences (p b 0.05) in mean whereas those that possess the same letter are statistically similar.

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Fig. 5. Gelation time for C:G:C hydrogels with different crosslinker ratios and CNC incorporation. C:G:C hydrogels were prepared without CNCs (X) or with CNCs with no surface charge (O). The inversion method at 37 °C was used to determine in situ gelation time. Results are an average of four independent measurements.

3.4. Bioactivity of cellulose nanocrystal/dual-crosslinked chitosan hydrogels As CNC/P:G:C hydrogels were being designed for VCF repair, their impact on mesenchymal stem cell (MSC) health and osteogenesis was explored. For these biological experiments, crosslinking ratio (i.e. 5:1, 10:1, and 15:1) were varied, but 2.5:1 genipin:chitosan ratio and 0CNC incorporation were used for all formulations due to the better mechanical properties observed with these hydrogels. 0CNC/C:G:C hydrogels were included as biologically inert control formulations. MSC proliferation and viability due to the presence of different hydrogel formulations are shown Fig. 7 for which statistical analysis was conducted (Table S4 and Table S5). MSCs seeded on tissue cultured plastic (Ctrl) and incubated in growth media proliferated to seven times initial seeding cell number over 14 days of culture (Fig. 7a). In contrast MSCs exposed to any hydrogel formulation showed some decrease in cell count in the first day but otherwise remained constant throughout the rest of the 14 days of the experiment (Fig. 7a). At each time point MSCs given growth media alone were found to be at statistically significantly greater cell counts than all hydrogel formulations among which no differences were observed (Table S4). An MTS assay showed that MSCs exposed to any hydrogel formulations for up to 14 days were N50% metabolically active compared to Ctrl treatment (Fig. 7b). No statistically significant differences in viability were observed among almost any groups or time points (Table S4 and Table S5). Importantly neither ionic crosslinker type (i.e. phosphate or

carbonate) nor ratio were found to dramatically impact MSC proliferation or viability. MSC ALP activity and mineralization due to interactions with different hydrogel formulations are provided in Fig. 8 for which statistical analysis was carried out (Table S4 and Table S5). MSCs exposed to media alone (Ctrl) showed background levels of ALP expression which were statistically significantly lower than any hydrogel formulation at each time point (Fig. 8a and Table S4). Furthermore, MSCs cultured with 0CNC/P:G:C hydrogels had statistically higher ALP activity than those presented with 0CNC/P:G:C hydrogels at days 7 and 14 (Fig. 8a and Table S4). The highest phosphate:genipin ratio (i.e. 15:1) yielded quicker increased ALP activity seen at days 3 and 7 but the lower phosphate:genipin ratios (i.e. 5:1 and 10:1) caught up by day 14 (Fig. 8a, Table S4 and Table S5). MSC mineralization evaluation revealed similar trends but even starker contrasts between the formulations (Fig. 8b). By day 7 the presence of any hydrogel yielded an increase in cell-based mineral deposition over the Ctrl group (Fig. 8b and Table S4). 0CNC/P: G:C hydrogels induced statistically significant mineralization over background more quickly (i.e. by day 3) than 0CNC/C:G:C hydrogels (Fig. 8b and Table S4). Phosphate content was also found to influence the extent of MSC mineralization with the highest phosphate:genipin ratio (i.e. 15:1) having greater ALZ staining than the lowest phosphate:genipin ratio (i.e. 5:1) at day 14 (Fig. 8b and Table S4). Regardless of phosphate:genipin ratio all 0CNC/P:G:C hydrogels showed increased cell-based mineralization over time (Fig. 8b and Table S5). 4. Discussion

Table 4 Swelling ratio and mass loss for C:G:C hydrogels with different crosslinking ratios and CNC incorporation. Mass loss was measured by immersing hydrogels in PBS at 37 °C for 14 days and comparing the change in mass over time. The swelling ratio was determined as the ratio of wet weight to dry weight of hydrogels after immersion in PBS for 24 h at 37 °C. Results are an average of four independent measurements. C:G:C

No

Swelling ratio 5:0.312:1 5:1.25:1 5:2.5:1 10:0.312:1 10:1.25:1 10:2.5:1 15:0.312:1 15:1.25:1 15:2.5:1

0

CNC

29.2 ± 1.6 26.6 ± 0.6 21.6 ± 1.0 28.6 ± 6.2 23.6 ± 4.5 19.4 ± 1.4 26.6 ± 2.8 22.1 ± 3.7 17.8 ± 0.9

Mass loss (%) 62.5 ± 4.8 55.2 ± 11.5 53.4 ± 10.7 61.5 ± 7.8 59.9 ± 17.8 55.3 ± 19.4 65.0 ± 5.1 56.3 ± 4.0 54.2 ± 9.4

CNC

Swelling ratio 28.7 ± 1.3 25.8 ± 2.1 19.8 ± 1.4 25.3 ± 3.3 22.7 ± 3.3 18.0 ± 1.6 23.6 ± 2.3 21.3 ± 0.4 16.9 ± 1.8

Mass loss (%) 45.4 ± 2.1 42.8 ± 5.9 37.0 ± 3.4 40.5 ± 5.4 38.0 ± 10.6 35.5 ± 9.6 43.8 ± 5.9 39.7 ± 16.1 34.3 ± 12.9

VCFs are known to facilitate impingement of the spine and a significant loss of movement and function. The gold-standard treatment for managing VCFs is through a minimally invasive injection of PMMA to stabilize the vertebral body and relieve pain. In this report, we aimed to develop a hydrogel technology that can better address VCFs focusing on resolving some of the disadvantages of PMMA including its lack of bioactivity, non-resorbability, highly exothermic gelation, and tissue incongruent mechanical properties. Important properties of the chitosanbased hydrogels that should be taken in to account for this biomedical application are gelation time, swelling ratio, stability, compressive strength, and bioactivity. Ideally, the hydrogel should form quick enough after injection (b20 min) to limit surgery time and infection risk while filling the defect efficiently [52]. The swelling ratio of hydrogels are a critical property to consider as it not only affects the mechanical properties and the stability of the material, but can also decrease the chances of vertebral body collapse [53]. The rate of hydrogel dissociation should match

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Fig. 6. Compressive strength for C:G:C hydrogels with different crosslinking ratios and CNC incorporation. C:G:C hydrogels were prepared without CNCs (X) and with CNCs with no surface charge (O). Results are an average of four independent measurements. Statistical analysis of the differences between groups are available in supplementary information (Table S1, Table S2, and Table S3).

the kinetics of host bone regeneration. Erosion that is too rapid can lead to mechanical failure of the developing neotissue while material mass loss that is too slow can inhibit new bone formation [54]. Hydrogel compressive strength should optimally be similar to human vertebrae bone because if it's too high it can cause stress shielding and greater secondary

fracture risk and if it's too low it will not stabilize the fracture site [55]. Vertebral bone compressive strength for healthy adults ranges from 0.6 to 6 MPa depending on bone density [56,57]. As VCFs primarily occur in osteoporotic patients, new technologies should be designed to match the properties of the local tissue environment which have been

Fig. 7. Proliferation (a) and viability (b) of MSCs exposed to different hydrogel formulations. Cell count was measured by the Quanti-iT PicoGreen Assay over 14 days of cell culture in media with or without 0CNC/Y:G:C X:2.5:1 hydrogels with different ionic crosslinkers and crosslinking ratios (N = 4). Cell metabolism was indirectly evaluated by NAD(P)H activity using a MTS assay over 14 days of cell culture in media with or without 0CNC/Y:G:C X:2.5:1 hydrogels with different ionic crosslinkers and crosslinking ratios (N = 4). Statistical analysis between groups and time points are available in supplementary information (Table S4 and Table S5).

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Fig. 8. Alkaline phosphatase (ALP) activity (a) and cell-based mineralization (b) of MSCs treated with different hydrogels. ALP enzyme activity was analyzed by an ALP Assay over 14 days of cell culture in media with or without 0CNC/Y:G:C X:2.5:1 hydrogels with different ionic crosslinkers and crosslinking ratios (N = 4). Alizarin red (ALZ) staining was used as an indirect measure of mineralization over 14 days of cell culture in media with and without 0CNC/Y:G:C X:2.5:1 hydrogels with different ionic crosslinkers and crosslinking ratios (N = 4). ALZ content for matching acellular hydrogel formulation over the same incubation time was subtracted to determine cell-based mineralization. ALP activity and mineralization were normalized on a per cell basis and statistical analysis between group and time points are available in supplementary information (Table S4 and Table S5).

found to be 20%–35% weaker than healthy vertebral bone [58]. Since not all VCFs are identical, a mechanically tunable material which can be customized to best fit each case is highly desirable. Finally, hydrogels capable of enhancing bone regeneration would help address the underlying osteoporosis-induced bone homeostasis dysregulation which is the primary underlying cause of VCFs. A variety of bioactive and biodegradable cements and hydrogels have been developed for VCF repair. Cements like calcium phosphate cement (CPC) or calcium sulfate cement (CSC) could overcome some of the drawbacks of PMMA, but their uncontrollable degradation limits their application in spine-associated disorder [59]. Hydrogels can possess more controllable gelation and degradation kinetics, but are commonly too soft for load-bearing applications as their compressive strengths are often b10 kPa [60–62]. The strength of the hybrid hydrogel detailed in this work is its ideal mechanical properties not previously observed with similar systems. This research investigated the effect ionic crosslinking type and ratio, covalent crosslinking ratio, and cellulose nanocrystal incorporation and surface charge had on chitosan-based hydrogel materials and biological properties to aid in the creation of a novel VCF treatment system. For ionic crosslinking, the impact of crosslinker type (i.e. phosphate or carbonate) and crosslinking ratio (i.e. 5:1, 10:1, or 15:1) were evaluated. The results show that the use of phosphate or a higher ionic crosslinking ratio greatly decreased gelation time (Fig. 1 and Fig. 5) and moderately enhanced compressive strength (Fig. 2 and Fig. 6). Ionic crosslinking ratio was also found to have a modest impact on swelling ratio (Table 2

and Table 4). Interestingly, mass loss was found to be considerably lower for carbonate crosslinked hydrogels than phosphate crosslinked hydrogels (Table 2 and Table 4). According to Kumar et al., the change in the type and functionality of the crosslinking reagent can greatly alter hydrogel microstructure impacting a variety of different materials properties [63]. Phosphate can achieve a negative charge state up to three which can be distributed among its four oxygens whereas carbonate can only reach a maximum negative charge state of two spread among its three oxygens. The amount of charge-associated complexation is also influenced by the ionic crosslinker ratio. The regional charge density increase achieved through the use of phosphate and/or high crosslinking ratios explains the faster gelation kinetics, resistance to swelling, and greater mechanical strength in accordance with previously published results [53,64]. In contrast, carbonate-based ionic crosslinking facilitates better resistance to dissociation which is potentially due to better molecular organization of the hydrogel components [63]. In addition to ionic crosslinking, the effect covalent crosslinking ratio had on hydrogel materials properties was explored. Increasing this ratio from 0.312:1 to 2.5:1 resulted in significantly decreased time to gelation (Fig. 1, Fig. 3, and Fig. 5) and hydrogel swelling (Table 2, Table 3, and Table 4) as well as moderate enhancements in dissociation resistance (Table 2, Table 3, and Table 4) and compressive strength (Fig. 2, Fig. 4a, and Fig. 6). As higher concentration of genipin is expected to increase crosslinking density, these results align with prior data tying hydrogel ductility, swelling, and strength to the quantity of covalent

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crosslinks [28,64]. The presence of even more covalent crosslinker (i.e. ratios of 5:1 and 7.5:1) yielded even stronger hydrogels (Fig. 4a), though these were found to be stiffer (Fig. 4b), to swell less (Table 3), and undergo greater mass loss (Table 3) than hydrogels with lower ratios. The brittleness and minimal swelling associated with these highly covalently crosslinked hydrogels [40] make them less desirable for VCF repair application. Interestingly, ionic and covalent crosslinking had synergistic benefits on the materials properties of chitosan hydrogels yielding highly swollen but strong biomaterials in accordance with previous research [32,40]. While dual-crosslinking can be optimized to enhance the materials properties of chitosan hydrogels, the greatest compressive strength achieved was about 45 kPa (P:G:C 15:2.5:1 - Fig. 3 C:G:C and 5:2.5:1 Fig. 6) which is still well below vertebral bone. Prior work has shown that the inclusion of a filler such as CNCs can enhance the biomechanical performance of biomaterials-based scaffolds [65]. CNCs are also easily synthetically manipulated allowing for their surface charge to be readily tuned potentially impacting their association with other hydrogel components as well as overall materials crosslinking. The presence of surface charge can cause a better dispersion of the filler within the aqueous solution as a polar solvent leads to more homogenous distribution within the chitosan matrix. However, CNCs with no surface charge can lead to more aggregated structures which may alter their capacity to be evenly dispersed [66]. The inclusion of negatively-charged CNCs (−CNC) or neutral CNCs (0CNC) increased gelation time whereas positively-charged CNCs (+CNC) did not (Fig. 1). This behavior is likely attributable to the crosslinking inhibition caused by competing ionic and physical interactions caused by the presence of −CNC or 0CNC. In contrast, +CNC has the same charge as chitosan allowing for it to participate in additional, instead of competing, ionic crosslinking. The presence of any CNC formulation (−CNC, 0CNC, or +CNC) decreased swelling and retarded dissociation of the hydrogels with 0CNC inducing the most drastic effects (Table 2). Physical entanglement due to any CNC filler is expected to decrease water penetration into the hydrogel though this is partially mitigated by the added hydrophilicity found with charged CNCs (−CNC or +CNC). Interestingly, only the incorporation of 0CNC enhanced the compressive strength of dual-crosslinked chitosan hydrogels (Fig. 2). The high aspect ratio CNC chains are expected to align with the direction of force resulting in higher compressive strength [67]. Unfortunately, the electrostatic association between −CNC or +CNC with other hydrogel components may limit this reorientation leading to the minimal changes in material strength observed. The physical property modifications achieved with 0CNC were also found to be independent of ionic crosslinker type (Fig. 5, Table 3, and Fig. 6). With the capacity to generate CNC-reinforced, dual-crosslinked hydrogels that possess similar materials properties regardless of ionic crosslinker type, the influence phosphate incorporation has on hydrogel toxicity and osteoinductivity was investigated. While some immediate modest cell death is seen with all hydrogel formulations, cell counts were stable from day 1 through day 14 (Fig. 7a). This cellular response is quite common due to the effect hydrogel components [68–70] and mechanotransduction [71,72] can have on cells during their initial contact with biomaterials. The cells that survived were found to be quite viable throughout the entire study (Fig. 7b). Excitingly, MSCs possessed increased ALP activity (Fig. 8a) and facilitated enhanced mineralization (Fig. 8b) when exposed to hydrogels compared to cells receiving no stimulus with the most significant improvement in osteoinductivity achieved through the incorporation of phosphate crosslinking.

CNC surface charge) imparts individual effects on hydrogel properties, the confluence of these variables on strengthening or weakening each other's effects were noticeable. Among all hydrogels, co-crosslinking with phosphate (15:1) and genipin (2.5:1) as well as incorporating neutral CNCs yielded a hydrogel that swells appropriately (ratio N 10), resists dissociation (b35% mass loss after 14 days), possesses near osteoporotic vertebral bone compressive strength (~150–250 kPa), and facilitates stem cell osteogenesis. Though promising, mild cytotoxicity associated with the hydrogel needs to be addressed going forward. Forthcoming studies will focus on additional hydrogel modifications including the incorporation of rapidly decomposing ceramics to better achieve the therapeutic window of osteoinductive calcium and phosphate ions. Also, in vivo studies will be conducted exploring the regenerative impact of these hydrogels in pre-clinical VCF models. Overall, the results shown provide significant initial support for the clinical translatability of this biomaterials-based platform to address VCFs in osteoporotic patients. Acknowledgments The authors gratefully acknowledge support from the Coulter Translational Partnership Program at the University of Missouri and start-up funds provided by the University of Missouri. Appendix A. Supplementary data Supplementary data to this article can be found online at https://doi. org/10.1016/j.ijbiomac.2019.02.086. References [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]

5. Conclusion This research suggests a multi-component chitosan-based injectable hydrogel can be engineered that possesses the mechanical properties and bioactivity desired for VCF repair. Although each component (ionic crosslinker type and content, covalent crosslinker content, and

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