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. Author manuscript; available in PMC: 2020 Aug 10.
Published in final edited form as: Biotechnol J. 2020 Mar 16;15(8):e1900456. doi: 10.1002/biot.201900456

Synthesis of Injectable Shear-Thinning Biomaterials of Various Compositions of Gelatin and Synthetic Silicate Nanoplatelet

Chengbin Xue 1,2,3,4, Huifang Xie 5,6,7, James Eichenbaum 8,9, Yi Chen 10,11,12, Yonggang Wang 13,14,15, Floor W van den Dolder 16,17,18,19, Junmin Lee 20,21, KangJu Lee 22,23, Shiming Zhang 24,25, Wujin Sun 26,27, Amir Sheikhi 28,29,30,31, Samad Ahadian 32,33, Nureddin Ashammakhi 34,35,36, Mehmet R Dokmeci 37,38,39, Han-Jun Kim 40,41, Ali Khademhosseini 42,43,44,45
PMCID: PMC7415533  NIHMSID: NIHMS1583100  PMID: 32107862

Abstract

Injectable shear-thinning biomaterials (iSTBs) have great potential for in situ tissue regeneration through minimally invasive therapeutics. Previously, an iSTB was developed by combining gelatin with synthetic silicate nanoplatelets (SNPs) for potential application to hemostasis and endovascular embolization. Hence, iSTBs are synthesized by varying compositions of gelatin and SNPs to navigate their material, mechanical, rheological, and bioactive properties. All compositions (each component percentage; 1.5–4.5%/total solid ranges; 3–9%) tested are injectable through both 5 Fr general catheter and 2.4 Fr microcatheter by manual pressure. In the results, an increase in gelatin contents causes decrease in swellability, increase in freeze-dried hydrogel scaffold porosity, increase in degradability and injection force during iSTB fabrication. Meanwhile, the amount of SNPs in composite hydrogels is mainly required to decrease degradability and increase shear thinning properties of iSTB. Finally, in vitro and in vivo biocompatibility tests show that the 1.5–4.5% range gelatin–SNP iSTBs are not toxic to the cells and animals. All results demonstrate that the iSTB can be modulated with specific properties for unmet clinical needs. Understanding of mechanical and biological consequences of the changing gelatin–SNP ratios through this study will shed light on the biomedical applications of iSTB on specific diseases.

Keywords: gelatin, injectability, laponite, rheology, shear-thinning biomaterial

1. Introduction

Shear-thinning hydrogels are non-Newtonian materials that behave as viscous fluids under shear stress and recover solid-like properties upon elimination of the stress.[1] Due to these properties, injectable shear-thinning biomaterials (iSTBs) are attracting attention as a group of self-healing materials that allow for fluent infusion and local equilibrium after approaching the final application site. In clinical applications, iSTBs can be delivered into the body using a needle or a general/microcatheters by manual pressure.[2–4] Compared to the conventional endovascular therapeutic agents (e.g., coil, stent), iSTB has the advantages of being easy to apply (no special equipment required), applicable to various size/shape of the lesions, and detectable through imaging modalities without artifact.[5,6] However, iSTB may require different physical characteristics depending on the clinical situation. For example, a catheter-based treatment technique requires a low injection force because iSTB must be delivered through a catheter of up to 150 cm through manual force, whereas a 3D-printing-based treatment technique requires enhanced structural stability after printing. Therefore, to meet the various clinical requirements, it is important to analyze the change in mechanical properties through the changes of the components that make up the iSTB.

The physical properties of iSTBs could be modulated by a combination of several carbon-based, polymeric, and inorganic nanomaterials.[7–10] In particular, inorganic nanomaterials have been widely reported for biomedical applications such as sustained drug release,[11] pH-responsive drug delivery,[12] and photothermal therapy.[13] Among them, synthetic silicate nanoplatelets (SNPs) have been employed in pharmaceutical, cosmetic, and food industries as active ingredients or rheological modifiers due to their unique electrostatic characteristics (formation of “house of cards” structure), uniform particle size, and bioactive properties.[14,15] SNPs are highly charged disks of ≈1 nm thickness and 20–30 nm diameter.[15] SNPs can form self-assembled structures that can form and break dynamically when in an aqueous medium due to anisotropic surface charges (edges; positive charge, surfaces; negative charge).[16] SNPs could improve mechanical properties of other polymeric biomaterials in tissue growth,[15,17,18] and show superior loading capacity for drugs/growth factors due to specific charged surface or high surface-to-volume ratio.[19] Based on these biomedically promising material properties, there have been attempts to use SNP as one of the iSTB components for biomedical application.[8,16,17,20–22]

We previously developed clinically applicable iSTBs based on biocompatible polymer (gelatin) mixed with SNP (i.e., Laponite XLG) for hemostasis and endovascular embolization.[23,24] These iSTBs exhibited strong shear-thinning behavior as well as biocompatible properties ranging from blood coagulation to minimized inflammatory response. Others have extended this work to implement iSTBs as embolic agents,[25] functionalized scaffold,[26,27] 3D bioinks,[28] and drug delivery systems.[29] Despite various STBs used, no studies have been reported on changes in the mechanical/bioactive properties due to the compositional differences of SNP-containing composite hydrogels. This gap in knowledge necessitates further investigation of such composite hydrogels. Here, we describe the swelling, degradation, microstructural, mechanical, and rheological properties of various compositions of composite hydrogels consisting of gelatin and SNPs. In addition, we analyzed both in vitro and in vivo biocompatibility and degradability according to the change in the ratio of gelatin and SNPs. We expect that the understanding of mechanical and biological consequences by the changing gelatin–SNP ratios will shed light on the biomedical applications of iSTB on specific diseases.

2. Results and Discussion

2.1. Injectable Gelatin/SNPs Hydrogel Compositionsand Characterization

Gelatin is a biocompatible polymer approved by the Food and Drug Administration, which limits the adsorption of nonspecific proteins, improves hemolysis, and ultimately prolongs clotting time, demonstrating substantially improved hemocompatibility of iSTB.[24,30–32] The introduction of gelatin can change the homogeneity and transparency of composite hydrogels and significantly reduce their pH-responsive properties.[22] Here, we hypothesized that changing the composition of gelatin–SNP-based hydrogels could have a significant impact on swellability, degradability, and mechanical and rheological properties by modulating physical and mechanical interactions of the materials. Based on our preliminary experiments for adjusting the physical properties of various composite hydrogels, here we prepared them with different compositions, that is, gelatin and SNPs (from 1.5% to 4.5%) to test the above hypothesis (Figure 1A).

Figure 1.

Figure 1.

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Compositions and characterizations of composite hydrogels. A) A gross view of different compositions with the concentration range of gelatin and SNPs from 1.5% to 4.5%. B) The degradability of composite hydrogels. C) The swellability of composite hydrogels. Data were expressed as mean ± SD, four samples for each group and each time point. *, Compared to 1.5%; #, Compared to 3.0%; *, p < 0.05; **, p < 0.01; one-way ANOVA analysis. D) Scanning electron microscopy (SEM) images of cross-sections from freeze dried nine composite hydrogel scaffolds. E) Statistical analysis of the SEM images (* vs 1.5%, # vs 3.0%; *p < 0.05, **p < 0.01; one-way ANOVA analysis). Data are means ± SD (n = 4).

The physical characteristics of prepared composite hydrogels were varied with concentrations of gelatin and SNPs. We observed gross morphology of the gels immediately after the final mixing stage and found trends across composition gradients (Figure 1A). Composite hydrogels with increasing percentage of gelatin appeared more of a solid at room temperature. High gelatin content with low SNP content resulted in a heterogeneous, rough appearance across the surface of the gel. Low concentration of both SNPs and gelatin resulted in a hydrogel that flowed easily and acted as a homogeneous solution, and those with a higher percentage of SNPs tended to have a more stable structure with the homogeneous surface. The latter traits were observed when either gelatin or SNP content reached 3% or more. These results are in line with previously reported papers.[23,24] iSTB-like properties are shown when the SNP content of the total solid percentage exceeds half of the total solid percentage within 3–9%. All nine compositions gave way under some applied stress, and the nature of the hydrogels’ shape recovery after elimination of stress was clarified with rheology experiments. With these initial observations, we could find that the hydrogels’ physical characteristics could be altered by changing the composition.

As an initial reflection of the host response, it is crucial to clarify the swelling capacity and degradation of injectable hydrogels in hemostatic and embolizing agent development.[33,34] Effective swelling capacity in iSTBs could lead to rapid hemostasis due to blood aggregation. In addition, the release of drugs or solvents from the polymeric mixture is mainly controlled by hydrogel swelling degree and rate.[35] Therefore, we analyzed the difference of swelling capacity and degradation according to the compositional changes of composite hydrogels. To test the degradability of composite hydrogels in vitro, different compositions (total 200 mg) were incubated with human plasma at 37 °C for 48 h (Figure 1B). All composite hydrogels with different compositions showed various degrees of swelling in the plasma ranging from 2% to 20% within 1 h. Interestingly, composite hydrogels were degraded mainly depended on the amount of SNPs in the hydrogels. Lowest SNP groups (gelatin 1.5%/SNPs 1.5% and gelatin 3.0%/SNPs 1.5%) exhibited the fastest degradation among all groups, while more than 3% of SNPs showed a degradation ratio less than 20% even after 48 h. To further evaluate the effect of each component on the swellability, the weight of PBS-immersed swollen composite hydrogels were compared to the freeze-dried hydrogels (Figure 1C). Most of the different composite hydrogels showed various degrees of swelling in PBS ranging from 500% to 1000% within 1 h. Our data indicated that swellability of composite hydrogels were mostly dependent on the amount of gelatin. With the increase of gelatin percentage, swellability of composite hydrogels decreases. Previously, iSTB was shown to have increased stability in a physiological environment mediated by robust interactions between ions within the iSTB[36] and the ability of SNPs to physically strengthen a hydrogel nanostructure.[37,38] Moreover, increased physical interactions among nanostructure of gelatin with specifically charged surfaces of SNPs may contribute to the formation of reversible nanostructure.[39,40] Thus, as with these results, our results show that gelatin rather than the SNP affects the swelling properties of iSTB.

The microstructure of hydrogels, including porosity and pore size can affect cell attachment, migration, and proliferation.[41–43] Small pore size scaffolds have an increasing surface area which will provide more cell binding sites for proliferation.[42] However, small pore size also limits cell migration and proliferation, followed by cell aggregations in periphery, which will limit the exchange of nutrients and waste in the center of scaffolds.[43] Therefore, it is crucial to analyze pore properties of the hydrogel for the cell-laden hydrogel–based clinical application. To determine whether compositional changes affect the structure of the composite hydrogel, the microstructure was further investigated using SEM after freeze-drying (Figure 1D). The samples with 4.5% gelatin had a much porous network (Figure 1E). However, in accordance with previous work, no significant changes are shown by adding SNP.[36] Our results show that the structure, degradability, and swellability of gelatin–SNP composite hydrogels are more influenced by gelatin than SNPs. In addition, these characteristics were found to be prominent as the gelatin content increased (G 4.5% > G 1.5%).

2.2. Injectability and Rheological Characteristics of Composite Hydrogels

To answer the question of whether our composite hydrogels shared the characteristics of injectable embolic biomaterials, we assessed injectability and shear-thinning characteristics of various compositions of composite hydrogels. We observed that all compositions could be delivered via 5 Fr general catheter/3 cc syringe and 2.4 Fr microcatheter/1 cc syringe combinations. Relatively stable increasing-plateaued curves without clumping or irregular injection force fluctuations were observed regardless of the physical appearance of the composite hydrogels (Figure 2A). The injection force measurement showed higher levels of injection force obtained by 5 Fr Beacon general catheter/3 cc syringe combination than that obtained by 2.4 Fr Beacon microcatheter/1 cc syringe combination. The results from using both catheters showed a similar increasing pattern among various compositions of gelatin and SNPs. The plateau injection force of all compositions of composite hydrogels demonstrated their injectability through both catheters for biological and medical applications such as hemostasis and endovascular embolization (Figure 2B). Although clinically relevant main input is generally known to be lower than 20 N, the actual injectability depends not only on the force exerted on the syringe, but also on the ergonomic design of the syringe and catheter, which may affect the force applied in general.[44,45] To establish a threshold of applicable injection force, we compare to downward push finger strength gathered from a study of 150 subjects.[46] Here, they found a (mean ± SD) strength for downward thumb pushing of (184.1 ± 52.2) N for males and (135.2 ± 30.4) N for females aged 21–30. Though much higher than the force needed to deliver common medical solutions as contrast agent and saline solution by catheters (injection force: 1 N),[23] the force required to inject composite hydrogels (injection force: 20–30 N) is below 2 standard deviations of the mean of pushing strength for both males and females aged 21–30. Therefore, the composite hydrogels can be easily applied by hand when the total concentration is less than 9.0% (2.4 Fr microcatheter) and 7.5% w/v (5 Fr general catheter), respectively. Our results show that all compounds can be injected in various size catheter/syringe combinations, but more than a certain percentage of the total composition (7.5%) may require a higher than normal injection force for clinical applications.

Figure 2.

Figure 2.

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Injection force and rheological evaluation of composite hydrogels. A) The curves of injection force from different compositions of composite hydrogels using 2.4 Fr microcatheter and 5 Fr general catheter. B) Quantifications of injection force plateau values from different compositions of composite hydrogels (* vs 1.5%, # vs 3.0%; *p < 0.05, **p < 0.01; one-way ANOVA analysis). Data are means ± SD (n = 4). C) Shear stress, viscosity, and storage moduli (G′) from different compositions of composite hydrogels after repeated application of high strain (100% oscillatory strain) over time. The far-right column shows the dependence of shear stress, viscosity, and change in storage modulus between low and high shear rates with the composition of the material.

Rheological evaluation is a powerful tool to characterize the viscoelasticity of composite hydrogels.[47] Shear-thinning behavior features a self-assembled structure that can be dynamically formed and destroyed and is an important factorin 3D printing as well as endovascular delivery.[24,28,29,48] In addition, self-recovery characteristic (solidification of iSTBs after injection) is critical to remain localized.[49] We analyzed the shear thinning properties of the composite hydrogels by the change of gelatin/SNPs (Figure 2C). All compositions exhibited steadily declining viscosity and concave-up stress-strain curves on logarithmic scales, indicating shear-thinning behavior. In addition, the composite hydrogels recovered to their original modulus over several cycles, demonstrating these hydrogels possess the common characteristics of iSTBs. Initial shear stress, viscosity, and shear-thinning properties increased with increasing concentrations of gelatin and SNPs in all samples (Figure 2C). In particular, iSTBs containing at least 3% SNPs have been found to measure storage modulus of over 100 Pa. Previously 6% or 9% w/v total solid percentage nanocomposite hydrogels have been proven to regain solid-like behavior rapidly following high strain (100% oscillatory strain) that led to liquid-like behavior in the nanocomposite hydrogels.[24] Interestingly, in our study, the weight percentage of SNP, rather than the gelatin weight percentage or total solid percentage, had a significant effect on the change in shear strength and viscosity. These exceptional physical properties of the hydrogels could be due to the non-covalent surface interactions between gelatin and SNPs. On highly charged, anisotropic SNP surfaces, gelatin can be reversibly adsorbed and desorbed on SNPs, resulting in changes in hydrogel properties.[39,50,51] Based on our results, it was confirmed that the swellability and freeze-dried hydrogel scaffold porosity depend on gelatin rather than SNP, while the shear-thinning properties depend on SNP rather than gelatin. In addition, these characteristics were found to be prominent as the gelatin or SNP content increased (G 4.5% > G 1.5%, SNP 4.5% > SNP 1.5%). However, from the level exceeding 7.5% of the total solid percentage, an increase in the injection force affects injectability. Thus, the ratio of gelatin and SNP amount needs to be adjusted for specific clinical purpose.

2.3. In Vitro and In Vivo Biocompatibility Assessment of iSTBs

To assess the clinical application of iSTBs, bioactive effects of iSTB compositional changes were evaluated. We further selected three different composite hydrogels (gelatin/SNPs = 4.5%/1.5%, 3%/3%, and 1.5%/4.5%) to investigate both in vitro and in vivo biocompatibility. Similar to the results of iSTBs using SNP previously reported, all three compositions of iSTBs did not show cytotoxicity to the NIH/3T3 cells and cells were proliferated upon time up to 14 days (Figure 3A). Interestingly, the quantitative analysis showed that the cell proliferation was higher in gelatin 3%–SNPs 3% group, followed by gelatin 1.5%–SNPs 4.5% group (Figure 3B). These could be due to increased SNPs, which provide electrostatically driven protein adhesion sites to the cells that enable cell binding.[17,52] SNPs are known to degrade into non-toxic components (Na+, Mg2+, Si(OH)4, Li+) under physiological conditions, of which cations can mediate cell adhesion through increased integrin interactions.[53,54] Our results suggest that the SNP-driven enhanced cell adhesion and the porosity/swellability of gelatin can provide cell proliferation and microenvironment in the hydrogel scaffold, ultimately leading to improved cytocompatibility.

Figure 3.

Figure 3.

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In vitro and in vivo biocompatibility assessment of composite hydrogels. A) NIH/3T3 fibroblasts were seeded on hydrogels-coated slides followed by culturing for 14 days and intermittent staining with live/dead cell imaging assay. B) Quantitative analysis of cytocompatibility assay. Data are means ± SD (n = 5). C) Gross and histological analyses of 1 and f4 weeks after subcutaneous implantation. The scale bars for low, high magnification were 1 mm and 50 μm. Dotted line; border between iSTB (right) and surrounding tissue (left). D) Quantitative analysis of remaining iSTB and degradation rate (%). (* vs G 4.5%/SNP 1.5%, # vs G 3.0%/SNP 3.0%; *p < 0.05, **p < 0.01,##p < 0.01; one-way ANOVA analysis). Data are means ± SD (n = 5).

Finally, we evaluated subcutaneous implantation for the biocompatibility and degradation in vivo and monitored over a 28-day period (Figure 3C,D). On day 7 HE staining results, all the iSTB groups we tested (gelatin/SNPs = 4.5%/1.5%, 3%/3%, and 1.5%/4.5%) did not cause significant inflammation (Figure 3C). On day 28, the amount of iSTB and surrounding granulation tissues of all groups decreased. In particular, in gelatin 4.5%/SNP 1.5% iSTB groups, most of the iSTB was degraded at 4 weeks. The quantitative analysis showed that the iSTB residual area under the skin was significantly different among groups (Figure 3D). Previously reported in vivo studies of iSTB showed that gelatin 2.25%/SNP 6.75% iSTB remained undissolved in the subcutaneous area of the mouse for 4 weeks, and another report showed gelatin 1.5%/SNP 4.5% iSTB remained in pig blood vessel for 24 days.[23,24] Similar to the results of in vitro degradation (Figure 1B), iSTBs containing more SNPs were found to degrade more slowly. Taken together, our results indicate gelatin–SNP-based iSTB showed biocompatible and is potentially capable of injectable therapeutics. Still, further evaluation of the potential effect on bioactivity as well as a biological disease model in vitro and in vivo will be required for future clinical translation.

3. Conclusion

To further understand gelatin–SNPs-based composite hydrogels, we investigated the mechanical, rheological, and bioactive properties of composite hydrogels with the various components of gelatin and SNPs. In our results, the increase of gelatin ratio causes the decrease in swellability, increase in freeze-dried hydrogel scaffold porosity, increase in degradability and injection force in addition to the known biocompatible polymers in iSTB fabrication. On the other hand, the amount of SNPs in composite hydrogels are mainly required to decrease degradability and increase shear thinning properties of iSTB. In addition, these characteristics were found to be prominent as the gelatin or SNP content increased (G 4.5% > G 1.5%, SNP 4.5% > SN P1.5%). All the compositions used in these experiments are injectable through different sized catheters and are biocompatible. However, more than a certain percentage of the total composition (7.5%) may require a higher force than normal injection force for clinical applications. Gelatin–SNPs-based iSTBs are reported to display superior biocompatibility/biostability, body temperature extrudability, shear-thinning behavior, and rapid network recoverability. These attractive physicochemical properties are favorable for easy administration in vivo, and a gelatin–SNPs-based iSTB may hold great potential in drug delivery, endovascular embolization, tissue regeneration, bioprinting, and for other biomedical applications. Understanding of mechanical and biological consequences of the changing gelatin–SNP ratios through this study will shed light on the biomedical applications of iSTB on specific diseases.

4. Experimental Section

Preparation of Composite Hydrogels and Scanning Electron Microscope

The preparation process of SNP/gelatin-based composite hydrogels was as described in previous reports.[23,55] Briefly, 18% w/v gelatin solution (type A, G1890, Sigma, CA, USA) and 9% w/v SNPs (Laponite XLG, BYK Additives Ltd, TX, USA) were prepared separately by using Milli-Q water. Composite hydrogels with different concentrations of gelatin (1.5%, 3.0%, 4.5%) and SNPs (1.5%, 3.0%, 4.5%) were synthesized from the stock solution.[23,24] According to preliminary studies, we limited the maximum concentration of gelatin and SNP to 4.5% because iSTB was difficult to inject through a catheter if the total solids of the iSTB exceeded 9%. The microstructure of dried hybrid composite hydrogels with various compositions was assessed by a scanning electron microscope (SEM) (Hitachi S-4800 FESEM, Hitachi, Tokyo, Japan) under an accelerating voltage of 1 kV. All samples for SEM imaging were kept at −80 °C for 24 h, freeze-dried for 48 h, and covered with a thin layer of Au. ImageJ was used to find porosity and average pore diameter according to previously reported protocols.[56,57] The porosity percentage and average pore diameter of the dried composite hydrogels were quantified from four independent sets of images.

Degradability of Composite Hydrogels

The degradability test was performed according to previous protocols with minor modification.[58,59] Briefly, composite hydrogels (200 mg) were incubated with human plasma (Sigma), in 1.5-mL Eppendorf tubes (VWR, CA, USA) at 37°C. After incubation for 1, 4, 6, 24, 28, and 48 h, the plasma was replaced with fresh human plasma and then the composite hydrogels were weighed. The relative weight of composite hydrogels percentage was defined as DR = (Wr/W0) × 100%, where Wr and W0 were the weight of the remaining composite hydrogels at various time points and the weight of composite hydrogels at the initial state, respectively. The experiments were repeated at least four times, and all data were averaged over all replicates.

Swellability of Composite Hydrogels

The swellability test was performed according to previous protocols with minor modification.[60,61] To assess the swelling ability of the composite hydrogels, the percentage of swelling was determined using the follow equation. Swelling ratio (SR) = (Ws − Wd)/Wd × 100%, where Ws and Wd were the weight of swollen hydrogels and freeze-dried weight, respectively. Completely freeze-dried hydrogels were immersed for a certain time period in phosphate buffered saline (PBS). After a certain amount of incubation (0, 10, 20, 30, and 60 min), the samples were taken from the solution. Then, the surrounding fluids were quickly wiped with filter paper and the swollen hybrid composite hydrogels’ weights were measured immediately. The experiments were repeated at least five times, and all data were averaged over all replicates.

Injectability of Composite Hydrogels

To test the injection force, composite hydrogels were loaded into medallion syringes (1 cc [MSS011], 3 cc [MSS031]; Merit Medical Systems, UT, USA) and injected through medical catheters (5 Fr Beacon, 2.4 Fr Beacon; Cook Incorporated, IN, USA). Following the combination of syringe–catheter used in the clinic, each material was tested using a 1 cc syringe with a 2.4 Fr microcatheter and a 3 cc syringe with a 5 Fr catheter. The injection force was monitored by a mechanical tester (Instron 5943, Instron Int. Ltd., MA, USA) with a 100 N load cell and recorded by Bluehill version 3 software, using an extension rate of 1.67 mm s−1. The average injection force of the plateau was obtained by quadruple measurements of three identical compositions of composite hydrogels.

Rheological Test of Composite Hydrogels

Rheological properties of composite hydrogels were evaluated by a Rheometer (AR-G2, TA instruments protocol) according to previously described protocols with minor modifications.[23,24] Shear stress, viscosity, and storage moduli were measured with a parallel plate geometry. Before testing, all composite hydrogel samples were equilibrated at room temperature for 2 min. To prevent water evaporation, mineral oil was added around the plate after composite hydrogel samples were loaded on the plate. For all samples, oscillatory stress sweep was completed at 0.1–1000 Pa under a fixed oscillatory frequency 1 Hz, and oscillatory frequency sweep was achieved at 0.1–100 Hz under fixed oscillatory stress 10 Pa at 25 °C. The shear stress, viscosity, and storage moduli were recorded by Anton Paar Rheocompass software.

In Vitro Cytocompatibility Assessment

The NIH/3T3 cells (ATCC; CRL1658, VA, USA) were cultured in Dulbecco’s Modified Eagle Medium (DMEM; Gibco, CA, USA) supplemented with 10% fetal bovine serum (Gibco), 50 μg mL−1 streptomycin, and 50 U mL−1 penicillin (Gibco) in 5% CO2 at 37 °C. NIH/3T3 cells were seeded in three different compositions of iSTB (gelatin 1.5%–SNP 4.5%, gelatin 3.0%–SNP 3%, gelatin 4.5%–SNP 1.5%) coated slide glass at a concentration of 1 × 105 cell and cultured for 1, 3, 7, and 14 days. After incubation for a certain period of time, cell viability was determined using live/dead viability/cytotoxicity kit (Invitrogen). Stained slides were imaged using fluorescence microscopy (Zeiss Axio Observer; Carl Zeiss, Jena, Germany). For the quantitative data analysis, five non-overlapping areas were counted at 200× magnification using ImageJ software (NIH, MD, USA). Viability (%) expressed the ratio of living cells to total cell numbers as mean ± SD.

In Vivo Biocompatibility Assessment

All animal experiments were approved by the UCLA Animal Research Committee (UCLA ARC #2017–096-01). The animal experiments were conducted aligned with relevant guidelines. 7-week-old, Sprague-Dawley male rats (average weight: 250–300 g) were bought from Charles River Laboratories (South San Francisco, CA, USA) and housed in approved animal facility. For the in vivo study, same three compositions of iSTBs that were tested in in vitro cytocompatibility assay were used. 2 cm incision was made on the posterior dorsal skin and iSTBs were implanted to the rats under inhalation anesthesia (1.5% isoflurane in 100% O2). Two subcutaneous pockets were made in each side of the incision and one iSTB per subcutaneous pocket was implanted (total of four experimental constructs per animal) for 1 and 4 weeks. After implantation, the incision was closed with sutures. To analyze the response of the host skin tissue to the application of the iSTB, animals were sacrificed using CO2. iSTB-implanted skin tissues were immediately collected and fixed in 10% neutral buffered formalin (Leica Biosystems, IL, USA). Fixed tissues were processed using standard methods and embedded in paraffin. 4 μm thickness skin tissue sections were stained with hematoxylin (Leica Biosystems) and eosin (Sigma) (HE) staining. Histology images were acquired on Nikon inverted microscope. Quantitative data such as residual iSTB area percentage, and degradation rate percentage were measured using the AmScope image analysis software (AmScope, Irvine, CA, USA).

Statistical Analysis

Data were displayed as mean ± standard deviation (SD). All statistical analyses and graphs were carried out by SPSS Statistics software (IBM, IL, USA) and GraphPad Prism 8.0 (GraphPad Software, CA, USA). Multiple comparisons were analyzed using one-way ANOVA with Tukey post hoc tests for more than triplicate of group data sets. p < 0.05 was classified as statistically significant for the tests.

Acknowledgements

C.X. and H.X. contributed equally to this work. The authors acknowledge funding from the National Institutes of Health (HL140951, HL137193). A.S. would like to acknowledge the postdoctoral fellowship from the Canadian Institutes of Health Research (CIHR) and the startup fund from the Pennsylvania State University.

Footnotes

Conflict of Interest

Ali Khademhosseini recently launched a start-up (Obsidiomed.com) based on a gel embolic for vascular embolization.

Contributor Information

Chengbin Xue, Center for Minimally Invasive Therapeutics (C-MIT), University of California, Los Angeles, Los Angeles, CA 90095, USA; Department of Bioengineering, Henry Samueli School of Engineering and Applied Sciences, University of California, Los Angeles, Los Angeles, CA 90095, USA; Key Laboratory of Neuroregeneration, Ministry of Education and Jiangsu Province, Co-Innovation Center of Neuroregeneration, Nantong University, Nantong, Jiangsu 226001, P. R. China; Jiangsu Clinical Medicine Center of Tissue Engineering and Nerve Injury Repair, Research Center of Clinical Medicine, Affiliated Hospital of Nantong University, Nantong, Jiangsu 226001, P. R. China.

Huifang Xie, Center for Minimally Invasive Therapeutics (C-MIT), University of California, Los Angeles, Los Angeles, CA 90095, USA; Department of Bioengineering, Henry Samueli School of Engineering and Applied Sciences, University of California, Los Angeles, Los Angeles, CA 90095, USA; Carbohydrate Laboratory, School of Food Science and Engineering, South China University of Technology, Guangzhou 510641 P. R. China.

James Eichenbaum, Center for Minimally Invasive Therapeutics (C-MIT), University of California, Los Angeles, Los Angeles, CA 90095, USA; Department of Bioengineering, Henry Samueli School of Engineering and Applied Sciences, University of California, Los Angeles, Los Angeles, CA 90095, USA.

Yi Chen, Center for Minimally Invasive Therapeutics (C-MIT), University of California, Los Angeles, Los Angeles, CA 90095, USA; Department of Bioengineering, Henry Samueli School of Engineering and Applied Sciences, University of California, Los Angeles, Los Angeles, CA 90095, USA; Beijing Biosis Healing Biological Technology Co., Ltd, Beijing 102600, P. R. China.

Yonggang Wang, Center for Minimally Invasive Therapeutics (C-MIT), University of California, Los Angeles, Los Angeles, CA 90095, USA; Department of Bioengineering, Henry Samueli School of Engineering and Applied Sciences, University of California, Los Angeles, Los Angeles, CA 90095, USA; Guangdong Engineering & Technology Research Center for Quality and Efficacy Reevaluation of Post-Market Traditional Chinese Medicine, Guangdong Key Laboratory of Plant Resources, School of Life Sciences, Sun Yat-sen University, Guangzhou 510275, China.

Floor W. van den Dolder, Center for Minimally Invasive Therapeutics (C-MIT), University of California, Los Angeles, Los Angeles, CA 90095, USA Department of Bioengineering, Henry Samueli School of Engineering and Applied Sciences, University of California, Los Angeles, Los Angeles, CA 90095, USA; Division of Heart and Lungs, Department of Cardiothoracic Surgery, University Medical Center Utrecht, Utrecht, GA 3508, The Netherlands; Regenerative Medicine Center Utrecht, University Medical Center Utrecht, Utrecht, CT 3584, The Netherlands.

Junmin Lee, Center for Minimally Invasive Therapeutics (C-MIT), University of California, Los Angeles, Los Angeles, CA 90095, USA; Department of Bioengineering, Henry Samueli School of Engineering and Applied Sciences, University of California, Los Angeles, Los Angeles, CA 90095, USA.

KangJu Lee, Center for Minimally Invasive Therapeutics (C-MIT), University of California, Los Angeles, Los Angeles, CA 90095, USA; Department of Bioengineering, Henry Samueli School of Engineering and Applied Sciences, University of California, Los Angeles, Los Angeles, CA 90095, USA.

Shiming Zhang, Center for Minimally Invasive Therapeutics (C-MIT), University of California, Los Angeles, Los Angeles, CA 90095, USA; Department of Bioengineering, Henry Samueli School of Engineering and Applied Sciences, University of California, Los Angeles, Los Angeles, CA 90095, USA.

Wujin Sun, Center for Minimally Invasive Therapeutics (C-MIT), University of California, Los Angeles, Los Angeles, CA 90095, USA; Department of Bioengineering, Henry Samueli School of Engineering and Applied Sciences, University of California, Los Angeles, Los Angeles, CA 90095, USA.

Amir Sheikhi, Center for Minimally Invasive Therapeutics (C-MIT), University of California, Los Angeles, Los Angeles, CA 90095, USA; Department of Bioengineering, Henry Samueli School of Engineering and Applied Sciences, University of California, Los Angeles, Los Angeles, CA 90095, USA; Department of Chemical Engineering, The Pennsylvania State University, University Park, PA 16802, USA; Department of Biomedical Engineering, The Pennsylvania State University, University Park, PA 16802, USA.

Samad Ahadian, Center for Minimally Invasive Therapeutics (C-MIT), University of California, Los Angeles, Los Angeles, CA 90095, USA; Department of Bioengineering, Henry Samueli School of Engineering and Applied Sciences, University of California, Los Angeles, Los Angeles, CA 90095, USA.

Nureddin Ashammakhi, Center for Minimally Invasive Therapeutics (C-MIT), University of California, Los Angeles, Los Angeles, CA 90095, USA; Department of Bioengineering, Henry Samueli School of Engineering and Applied Sciences, University of California, Los Angeles, Los Angeles, CA 90095, USA; Department of Radiological Sciences, David Geffen School of Medicine, University of California–Los Angeles, Los Angeles, CA 90095, USA.

Mehmet R. Dokmeci, Center for Minimally Invasive Therapeutics (C-MIT), University of California, Los Angeles, Los Angeles, CA 90095, USA Department of Bioengineering, Henry Samueli School of Engineering and Applied Sciences, University of California, Los Angeles, Los Angeles, CA 90095, USA; Department of Radiological Sciences, David Geffen School of Medicine, University of California–Los Angeles, Los Angeles, CA 90095, USA.

Han-Jun Kim, Center for Minimally Invasive Therapeutics (C-MIT), University of California, Los Angeles, Los Angeles, CA 90095, USA; Department of Bioengineering, Henry Samueli School of Engineering and Applied Sciences, University of California, Los Angeles, Los Angeles, CA 90095, USA.

Ali Khademhosseini, Center for Minimally Invasive Therapeutics (C-MIT), University of California, Los Angeles, Los Angeles, CA 90095, USA; Department of Bioengineering, Henry Samueli School of Engineering and Applied Sciences, University of California, Los Angeles, Los Angeles, CA 90095, USA; Department of Radiological Sciences, David Geffen School of Medicine, University of California–Los Angeles, Los Angeles, CA 90095, USA; Department of Chemical and Biomolecular Engineering, Henry Samueli School of Engineering and Applied Sciences, University of California–Los Angeles, Los Angeles, CA 90095, USA.

References

  • [1].Guvendiren M, Lu HD, Burdick JA, Soft Matter 2012, 8, 260. [Google Scholar]
  • [2].Purcell BP, Lobb D, Charati MB, Dorsey SM, Wade RJ, Zellars KN, Doviak H, Pettaway S, Logdon CB, Shuman JA, Freels PD, Gorman JH 3rd, Gorman RC, Spinale FG, Burdick JA, Nat. Mater. 2014, 13, 653. [DOI] [PMC free article] [PubMed] [Google Scholar]
  • [3].Jahan K, Mekhail M, Tabrizian M, Carbohydr. Polym. 2019, 203, 60. [DOI] [PubMed] [Google Scholar]
  • [4].Yang JA, Yeom J, Hwang BW, Hoffman AS, Hahn SK, Prog. Polym. Sci. 2014, 39, 1973. [Google Scholar]
  • [5].Hoffman AS, Adv. Drug Delivery Rev. 2002, 54, 3. [DOI] [PubMed] [Google Scholar]
  • [6].Weng L, Rostambeigi N, Zantek ND, Rostamzadeh P, Bravo M, Carey J, Golzarian J, Acta Biomater. 2013, 9, 8182. [DOI] [PubMed] [Google Scholar]
  • [7].Paul A, Hasan A, Kindi HA, Gaharwar AK, Rao VT, Nikkhah M, Shin SR, Krafft D, Dokmeci MR, Shum-Tim D, Khademhosseini A, ACS Nano 2014, 8, 8050. [DOI] [PMC free article] [PubMed] [Google Scholar]
  • [8].Su D, Jiang L, Chen X, Dong J, Shao Z, ACS Appl. Mater. Interfaces 2016, 8, 9619. [DOI] [PubMed] [Google Scholar]
  • [9].Mulyasasmita W, Cai L, Hori Y, Heilshorn SC, Tissue Eng., Part A 2014, 20, 2102. [DOI] [PMC free article] [PubMed] [Google Scholar]
  • [10].Whitlow J, Pacelli S, Paul A, Macromol. Chem. Phys. 2016, 217, 1245. [DOI] [PMC free article] [PubMed] [Google Scholar]
  • [11].Liao YT, Lee CH, Chen ST, Lai JY, Wu KCW, J. Mater. Chem. B 2017, 5, 7008. [DOI] [PubMed] [Google Scholar]
  • [12].Liang Y-H, Liu C-H, Liao S-H, Lin Y-Y, Tang H-W, Liu S-Y, Lai IR, Wu KCW, ACS Appl. Mater. Interfaces 2012, 4, 6720. [DOI] [PubMed] [Google Scholar]
  • [13].Liao YT, Liu CH, Chin Y, Chen SY, Liu SH, Hsu YC, Wu KCW, J. Mater. Chem. B 2019, 7, 4451. [Google Scholar]
  • [14].Dawson JI, Oreffo ROC, Adv. Mater. 2013, 25, 4069. [DOI] [PubMed] [Google Scholar]
  • [15].Gaharwar AK, Mihaila SM, Swami A, Patel A, Sant S, Reis RL, Marques AP, Gomes ME, Khademhosseini A, Adv. Mater. 2013, 25, 3329. [DOI] [PubMed] [Google Scholar]
  • [16].Ruzicka B, Zaccarelli E, Soft Matter 2011, 7, 1268. [Google Scholar]
  • [17].Tomas H, Alves CS, Rodrigues J, Nanomedicine 2018, 14, 2407. [DOI] [PubMed] [Google Scholar]
  • [18].Dawson JI, Kanczler JM, Yang XBB, Attard GS, Oreffo ROC, Adv. Mater. 2011, 23, 3304. [DOI] [PubMed] [Google Scholar]
  • [19].Gaharwar AK, Cross LM, Peak CW, Gold K, Carrow JK, Brokesh A, Singh KA, Adv. Mater. 2019, 31, 1900332. [DOI] [PMC free article] [PubMed] [Google Scholar]
  • [20].Li YL, Maciel D, Tomas H, Rodrigues J, Ma H, Shi XY, Soft Matter 2011, 7, 6231. [Google Scholar]
  • [21].Yang H, Hua S, Wang W, Wang A, Iran. Polym. J. 2011, 20, 479. [Google Scholar]
  • [22].Li C, Mu C, Lin W, Ngai T, ACS Appl. Mater. Interfaces 2015, 7, 18732. [DOI] [PubMed] [Google Scholar]
  • [23].Avery RK, Albadawi H, Akbari M, Zhang YS, Duggan MJ, Sahani DV, Olsen BD, Khademhosseini A, Oklu R, Sci. Transl. Med. 2016, 8, 365ra156. [DOI] [PubMed] [Google Scholar]
  • [24].Gaharwar AK, Avery RK, Assmann A, Paul A, McKinley GH, Khademhosseini A, Olsen BD, ACS Nano 2014, 8, 9833. [DOI] [PMC free article] [PubMed] [Google Scholar]
  • [25].Moore MJ, Malaxos L, Doyle BJ, J. Mech. Behav. Biomed. Mater. 2019, 99, 66. [DOI] [PubMed] [Google Scholar]
  • [26].Kerativitayanan P, Tatullo M, Khariton M, Joshi P, Perniconi B, Gaharwar AK, ACS Biomater. Sci. Eng. 2017, 3, 590. [DOI] [PubMed] [Google Scholar]
  • [27].Hasany M, Thakur A, Taebnia N, Kadumudi FB, Shahbazi M-A, Pierchala MK, Mohanty S, Orive G, Andresen TL, Foldager CB, Yaghmaei S, Arpanaei A, Gaharwar AK, Mehrali M, Dolatshahi-Pirouz A, ACS Appl. Mater. Interfaces 2018, 10, 34924. [DOI] [PubMed] [Google Scholar]
  • [28].Wilson SA, Cross LM, Peak CW, Gaharwar AK, ACS Appl. Mater. Interfaces 2017, 9, 43449. [DOI] [PubMed] [Google Scholar]
  • [29].Gharaie SS, Dabiri SMH, Akbari M, Polymers 2018, 10, 1317. [DOI] [PMC free article] [PubMed] [Google Scholar]
  • [30].Higuchi A, Kumar SS, Benelli G, Ling Q-D, Li H-F, Alarfaj AA, Munusamy MA, Sung T-C, Chang Y, Murugan K, Prog. Mater. Sci. 2019, 103, 374. [Google Scholar]
  • [31].Hardy JG, Lee JY, Schmidt CE, Curr. Opin. Biotechnol. 2013, 24, 847. [DOI] [PubMed] [Google Scholar]
  • [32].Park H, Temenoff JS, Tabata Y, Caplan AI, Mikos AG, Biomaterials 2007, 28, 3217. [DOI] [PMC free article] [PubMed] [Google Scholar]
  • [33].Mirzakhanian Z, Faghihi K, Barati A, Momeni HR, J. Biomater. Sci., Polym. Ed. 2015, 26, 1439. [DOI] [PubMed] [Google Scholar]
  • [34].Behrens AM, Sikorski MJ, Li T, Wu ZJ, Griffith BP, Kofinas P, Acta Biomater. 2014, 10, 701. [DOI] [PubMed] [Google Scholar]
  • [35].Ganji F, Vasheghani-Farahani S, Vasheghani-Farahani E, Iran. Polym. J. 2010, 19, 375. [Google Scholar]
  • [36].Lokhande G, Carrow JK, Thakur T, Xavier JR, Parani M, Bayless KJ, Gaharwar AK, Acta Biomater. 2018, 70, 35. [DOI] [PMC free article] [PubMed] [Google Scholar]
  • [37].Xavier JR, Thakur T, Desai P, Jaiswal MK, Sears N, Cosgriff-Hernandez E, Kaunas R, Gaharwar AK, ACS Nano 2015, 9, 3109. [DOI] [PubMed] [Google Scholar]
  • [38].Gaharwar AK, Kishore V, Rivera C, Bullock W, Wu C-J, Akkus O, Schmidt G, Macromol. Biosci. 2012, 12, 779. [DOI] [PubMed] [Google Scholar]
  • [39].Schexnailder P, Schmidt G, Colloid Polym. Sci. 2009, 287, 1. [Google Scholar]
  • [40].Chimene D, Alge DL, Gaharwar AK, Adv. Mater. 2015, 27, 7261. [DOI] [PubMed] [Google Scholar]
  • [41].Yuan L, Li XY, Ge LM, Jia XQ, Lei JF, Mu CD, Li DF, ACS Appl. Mater. Interfaces 2019, 11, 269. [DOI] [PubMed] [Google Scholar]
  • [42].O’Brien FJ, Harley BA, Yannas IV, Gibson LJ, Biomaterials 2005, 26, 433. [DOI] [PubMed] [Google Scholar]
  • [43].Murphy CM, Haugh MG, O’Brien FJ, Biomaterials 2010, 31, 461. [DOI] [PubMed] [Google Scholar]
  • [44].Öhrlund Å, Edsman K, Sturesson C, Nord L, Kenne A, Näsström J, 8th Anti-Aging Medicine World Congress, EuroMediCom, Paris: 2010, pp. 8–10. [Google Scholar]
  • [45].Vo A, Doumit M, Rockwell G, J. Med. Eng. 2016, 2016, 5162394. [DOI] [PMC free article] [PubMed] [Google Scholar]
  • [46].Peebles L, Norris B, Appl. Ergon. 2003, 34, 73. [DOI] [PubMed] [Google Scholar]
  • [47].Wang H, Wu Y, Cui C, Yang J, Liu W, Adv. Sci. 2018, 5, 1800711. [DOI] [PMC free article] [PubMed] [Google Scholar]
  • [48].Peak CW, Stein J, Gold KA, Gaharwar AK, Langmuir 2018, 34, 917. [DOI] [PubMed] [Google Scholar]
  • [49].Jalili NA, Jaiswal MK, Peak CW, Cross LM, Gaharwar AK, Nanoscale 2017, 9, 15379. [DOI] [PMC free article] [PubMed] [Google Scholar]
  • [50].Gaharwar AK, Peppas NA, Khademhosseini A, Biotechnol. Bioeng. 2014, 111, 441. [DOI] [PMC free article] [PubMed] [Google Scholar]
  • [51].Saikia J, Yazdimamaghani M, Moghaddam SPH, Ghandehari H, ACS Appl. Mater. Interfaces 2016, 8, 34820. [DOI] [PMC free article] [PubMed] [Google Scholar]
  • [52].Khalili AA, Ahmad MR, Int. J. Mol. Sci. 2015, 16, 18149. [DOI] [PMC free article] [PubMed] [Google Scholar]
  • [53].Zreiqat H, Howlett C, Zannettino A, Evans P, Schulze-Tanzil G, Knabe C, Shakibaei M, J. Biomed. Mater. Res. 2002, 62, 175. [DOI] [PubMed] [Google Scholar]
  • [54].Thompson DW, Butterworth JT, J. Colloid Interface Sci. 1992, 151, 236. [Google Scholar]
  • [55].Sheikhi A, Afewerki S, Oklu R, Gaharwar AK, Khademhosseini A, Biomater. Sci. 2018, 6, 2073. [DOI] [PMC free article] [PubMed] [Google Scholar]
  • [56].Anguiano M, Castilla C, Maska M, Ederra C, Pelaez R, Morales X, Munoz-Arrieta G, Mujika M, Kozubek M, Munoz-Barrutia A, Rouzaut A, Arana S, Garcia-Aznar JM, Ortiz-de-Solorzano C, PLoS One 2017, 12, e0171417. [DOI] [PMC free article] [PubMed] [Google Scholar]
  • [57].Xiao W, Liu W, Sun J, Dan X, Wei D, Fan H, J. Bioact. Compat. Polym. 2012, 27, 327. [Google Scholar]
  • [58].Ren Y, Zhao X, Liang X, Ma PX, Guo B, Int. J. Biol. Macromol. 2017, 105, 1079. [DOI] [PubMed] [Google Scholar]
  • [59].Assmann A, Vegh A, Ghasemi-Rad M, Bagherifard S, Cheng G, Sani ES, Ruiz-Esparza GU, Noshadi I, Lassaletta AD, Gangadharan S, Tamayol A, Khademhosseini A, Annabi N, Biomaterials 2017, 140, 115. [DOI] [PMC free article] [PubMed] [Google Scholar]
  • [60].Liu X, Chen W, Gustafson CT, Miller Ii AL, Waletzki BE, Yaszemski MJ, Lu L, RSC Adv. 2015, 5, 100824. [DOI] [PMC free article] [PubMed] [Google Scholar]
  • [61].Akakuru O, Isiuku B, J. Phys. Chem. Biophys. 2017, 7, 2161. [Google Scholar]

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