Microfluidic Manufacturing of Alginate Fibers with Encapsulated

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Microfluidic Manufacturing of Alginate Fibers with Encapsulated Astrocyte Cells Marilyn C McNamara, Farrokh Sharifi, Jasmin Okuzono, Reza Montazami, and Nicole N Hashemi ACS Appl. Bio Mater., Just Accepted Manuscript • DOI: 10.1021/acsabm.9b00022 • Publication Date (Web): 18 Mar 2019 Downloaded from http://pubs.acs.org on March 19, 2019

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ACS Applied Bio Materials

Microfluidic Manufacturing of Alginate Fibers with Encapsulated Astrocyte Cells Marilyn C. McNamara 1, Farrokh Sharifi 1, Jasmin Okuzono 2, Reza Montazami 1, Nicole N. Hashemi 1,3* 1

Department of Mechanical Engineering, Iowa State University, Ames, IA 50011

2

Department of Biological and Chemical Engineering, Iowa State University, Ames, IA

50011 3

Department of Biomedical Sciences, Iowa State University, Ames, IA 50011

* E-mail: [email protected]

Abstract Encapsulating cells within microfibers allows for immobilization with a high degree of spatial-temporal control. Furthermore, microfluidic encapsulation allows for the continuous creation of tunable fibers using mild, cell-friendly gelation conditions, making it advantageous over other fabrication methods. Mouse astrocyte cells (MACs) encapsulated within microfluidically-produced alginate fibers had a 24-hour survival rate of up to 89%, with up to 60% of cells surviving 11 days of encapsulation. The Young’s Modulus of both dry and wet fibers were found to be within the range of 400 and 17,000 MPa for dry fibers, and 20 and 90 MPa for wet fibers and wet cell-encapsulated fibers. Porosities between 12% and 92% were achieved. Keywords: astrocytes, cell encapsulation, microfluidics, hydrogel, microfiber

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1. Introduction While traditional 2-dimensional (2D) cell culturing techniques have provided crucial advancements for the field of medicine, they fail to recapitulate vital mechanical, spatial and chemical conditions cells face within their natural microenvironments.1-5 Not only can 3-dimensional (3D) cell culturing provide additional levels of complexity and physiological relevancy, it also offers potential for scalability and reduced cost.6-7 These approaches call upon the use of hydrogels to create spatially organized structures that may support and facilitate cell viability and growth.2, 7-10 Hydrogels are particularly well suited for cell culturing due to their diffusivity, which provides ingress for oxygen and nutrients, and egress for cellular waste products.11-12 Encapsulating cells within hydrogels is known to protect them from unnatural hydrodynamic stresses which arise when plating them on flat surfaces, and therefore can reduce uncontrolled differentiation of stem cells while also controlling aggregate size and supporting pluripotency for long-term trials.13-16 Additionally, hydrogels can be manipulated to cause the controlled release of therapeutic proteins and other beneficial chemicals aiding in the health and longevity of cell cultures.17-18 Due to their tissue-like water content and biocompatible nature, hydrogels are commonly used for 3D cell culturing,15,

19-21

thereby providing highly tunable and relevant

microenvironments to encourage cell growth and proliferation.2-3,

9, 22

This is a crucial

factor in 3D cell culturing applications, where even the smallest changes greatly affect the health and physiological relevancy of the resulting model.23 The goal is to reliably mimic the physical, mechanical and chemical microenvironments that cells experience in their native situations. Although multiple techniques can supply support systems for 3D cell culturing, cell-encapsulated microfibers provide both precise control over cell location and an increased surface-to-volume ratio over traditional film-based scaffolding techniques, which provides additional surface area for diffusion-based transport of nutrients to cell.14 Astrocytes were used for this encapsulation because of their prevalence in the central nervous system (CNS), where their quantities are five times greater than that of neurons.24 They are known to help regulate blood flow in the brain, while aiding in maintaining molecular homeostasis and contributing to CNS metabolism and synaptic


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functioning.24 During CNS injury, astrocytes have been shown to regulate the amount of CNS inflammation, affect blood-brain barrier permeability, and aid in tissue repair and synaptic remodeling; their pathologies are increasingly thought to cause or contribute to neuropathologies.25-26 For these reasons, astrocytes are used in models to help understand the repercussions and secondary damages resulting from traumatic brain injuries (TBIs) and other diseases;27-30 therefore, creating a relevant 3D astrocyte-based cell culture model will enable further research into TBI and neuropathology conditions. Polymerized alginate forms a naturally-derived hydrogel which is commonly used for biomedical engineering.22, 31-33 Alginate is a polysaccharide which is formed by blocks of 1-4 linked β-D-mannuronic acid (M block) and α-L-guluronic acid (G block).11-12,

19, 34

Alginate is polymerized via chemical crosslinking when divalent ions such as Ca2+, poly(ethylene glycol)-diamines, and methyl ester

L-lysine

link guluronic residues.11

Although it is highly biocompatible, the mechanical properties degrade rapidly when alginate-based hydrogels are in contact with physiological levels of sodium chloride (NaCl).34 However, a method of fiber fabrication with a high degree of control over the fiber size, shape, and mechanical properties can help to counteract the negative effects of NaCl-caused degradation. Microfluidic devices are versatile technologies, and microfluidic spinning is capable of creating highly controllable, continuously fabricated, cell-embedded fibers.7-9, 18, 35-40 This efficient, rapid, and low-cost technique does not require large sample sizes or extreme polymerization requirements.7 It relies upon the use of microchannels as well as pre-gel and sheath solutions to help shape and solidify the resulting microfiber.1-2,

7, 9

Fiber

polymerization can occur via chemical cross-linking and photopolymerization;36 both are suitable for cell encapsulation, although the latter relies upon the use of ultra-violet light, which can be damaging to sensitive cells. For fibers created with ionic polymerization, solidification begins as soon as the alginate and ionic solutions come into contact within the microfluidic device and continues as the solutions travel through the rest of the channel. This early polymerization yields a surface topography that is not smooth, since the gel is affected by the vertical motion introduced by the chevrons in the channel of the device.



Parameters such as concentrations of solutions and flow rates of sheath and core solutions provide a range of microfibers with different characteristics. For instance, increasing the sheath fluid flow rate but holding the core fluid rate constant will result in smaller fibers; similarly, increasing the concentration of the ionic cross-linker in the water bath leads to a higher degree of polymerization, which corresponds to tighter bond lengths and smaller fibers.7-8, 22, 41 Another crucial parameter to consider is the Flow Rate Ratio (FRR), or the ratio of the sheath fluid flow rate to the core fluid flow rate. These factors affect the volume of liquid travelling through the channel at any given time. Increasing the FRR will lead to a decrease in fiber size, and thereby will affect the mechanical properties of the resulting fibers.8, 22 In this study, the tensile properties of microfibers were analyzed in order to ascertain the elastic (Young’s) Modulus values of the fibers, which gives information regarding a material’s stiffness, and is generated from the slope of the linear region of the stress-strain curve.42 Such mechanical properties are worth considering when designing materials for use in biocompatible structures, where mimicking the microenvironment of target tissues is of vital importance.22 This is of particular interest, since it allows for the study of a variety neuropathology situations where a change in brain stiffness has been observed, including Alzheimer’s,43 Dementia,44 TBI,45 and more.46-47 In this study, Type 1 mouse astrocyte cells (MACs) were encapsulated into fibers created of pure alginate using microfluidic fiber fabrication with a well-documented microfluidic channel.7-8 Fiber polymerization of alginate was accomplished via ionically-driven chemical crosslinking of alginate by the Ca2+ found in Calcium Chloride Dihydrate (CaCl2 · 2H2O). Due to the high viscosity of alginate, it was necessary to increase the viscosity of the sheath fluid; therefore, polyethylene glycol (PEG) was added to the CaCl2 · 2H2O solution. This minimized variation in shear strain across the interfaces, thereby reducing fiber deformation and minimizing clogging within the channel.48 The significance of this study resides in the manufacturing process used to create the fibers. While alginate hydrogels are well understood in other methods, their use in creating fibers and microfluidically encapsulating astrocytes or other neuronal cells is limited. Microfluidic fiber fabrication is well suited for scaffolding and therapeutic


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applications due to its high control of fiber shape, size, and mechanical properties, all of which can affect cell viability and behavior. Therefore, the methodology in this study provides a framework which can benefit many applications. Similar fibers have been used to direct neural cell growth or affect cellular differentiation, both of which are vital within the field of neural tissue engineering.8, 10 Likewise, fibers which are able to spatially restrict the locations of cells over time while protecting them from external stimuli are highly desirable to regenerative medicines and other targeted cell therapy procedures.22

2. Experimental Section Creation of Microfluidic Device and Cell Preparation The microfluidic mold used for this study was created on a Silicon wafer using soft photolithography; it is described further in depth by Sharifi et al.7-8 To create the microfluidic device, polydimethylsulfoxide (PDMS) (Dow Corning, Midland, MI) was mixed in a 1:10 ratio of elastomer curing agent to elastomer base and was poured onto the molds. After allowing time for degassing, the two halves were hardened at 80 °C for 25 minutes. Layers were added via plasma cleaning on medium strength for 20 seconds for added thickness. The two halves were prepared using the same plasma cleaning process, and then were lined up under a microscope. Devices were sterilized with a 70% Ethanol rinse and a minimum of five hours of exposure to ultra-violet light within a biological fume hood. Preparation of Solutions and Frames The pre-gel solution was made by dissolving 6% alginate (Very Low Viscosity, Alfa Aesar, Ward Hill, MA) in WFI-Quality Cell Culture Grade Water (Corning, Corning, NY) at room temperature overnight. Before use, pre-gel solutions were filtered using syringe filters under pressure in the following sequence: 3 µm pore size, 0.45 µm pore size, and sterile 0.22 µm pore size. For the sheath fluid, 0.5% Calcium Chloride Dihydrate (CaCl2 · 2H2O) (Fisher Chemical, Waltham, MA), 5% Poly-ethylene Glycol (PEG) (Mn = 20 000, Aldrich Chemistry, St. Louis, MO) solutions were created within deionized water. 5% CaCl2 · 2H2O and pure Cell-culture grade water solutions were used for the 5% and 0% CaCl2 · 2H2O baths, respectively.



Frames were created to store the fibers within wells, and keep them stationary for imaging. Copper alone caused cell death, so copper frames were created and then placed in uncured PDMS, which was then set in an oven at 80°C until hard. Frames were sterilized with 70% Ethanol and UV exposure before use. Cell Culture Procedures Mouse astrocyte cells (ATCC C8-D1A) were cultured with Advanced Dulbecco’s modified Eagle’s medium (DMEM)/F12 (1X) (Gibco, Dublin, Ireland), which was supplemented with 10 % FBS (One ShotTM format, ThermoFisher Scientific, Waltham, MA), and 1% penicillin (10 000 U/mL) – streptomycin (10 000 µg/mL) (FisherScientific, Waltham, MA). All cell protocols follow the methods suggested by the ATCC product sheet. Cells were labeled in vitro using CellTracker™ Green Fluorescent probe CMFDA (Invitrogen, Carlsbad, CA), which was purchased from Thermo Fisher Scientific. The CMFDA was used by diluting it to 25 µM in serum-free media. Dye media was introduced to the flask and was allowed to sit for 45 minutes, at which point the dye was removed and replaced with regular maintenance media. Cells were allowed to rest for two hours before they were released with Trypsin EDTA (1x) (Cascade Biologics, Portland, OR) solution. Loose cells were collected in media and were centrifuged at 2,000 rpm for 5 minutes in order to create a cell suspension with a cell density 1.395∙106 cell/mL (viable cells) in the final alginate/cell suspension/media solution. This was accomplished by mixing 6% alginate with the cell suspension/media solution in a ratio of 1:3. Cell Encapsulation within Alginate Fibers The resulting mixture was placed within a 3 mL syringe, and was injected into the microfluidic device at a set rate with a syringe pump. A sheath fluid of 0.5% CaCl2 · 2H2O and 5% PEG polymerized the fiber within the microchannel. After microfibers were created, half were exposed to a 5% CaCl2 · 2H2O solution bath in order to further strengthen the existing fiber, whereas the other half was fabricated within a bath of pure DI water. Fibers were wound around the frames described above and were submerged in standard media in a 24 well plate with 700 µL of maintenance media and were stored in an incubator at 37˚C and 5% CO2. Every two days an additional 300 µL was added to


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ensure enough nutrients for cell health. No media was removed in order to ensure fibers remained within the wells. It was observed that PBS degraded fiber integrity and affected their degradation rates and mechanical properties. Because of this, PBS was not suitable for use in a collection bath. Previous results showed that hydrogel fibers protected cells for brief exposure to pure, sterile cell culture water, so this substitute was chosen for a bath.

Live/Dead Cell Assay To complete the Live/Dead cell assay, existing media was removed from the wells and the fibers were rinsed once with 200 µL of FBS and Penn-Strep- free DMEM. The rinse media was removed and replaced with 200 µL of a solution of 5 µM CellTrackerTM CMFDA and 4 µM Propidium Iodide (PI) (Invitrogen, Carlsbad, CA) dissolved into plain DMEM and warmed to room temperature. After incubation for 25 minutes at 37° C and 5% CO2, the staining solution was removed, and the fibers were rinsed again with 200 µL of plain DMEM. Upon removal, another 200 µL of plain DMEM was introduced into the well to prevent samples from drying before imaging was complete. For all stages, removal of the media must be completed with utmost care, to avoid undesired removal of the samples. Microscopy Images were collected with an Axio Observer Z1 Inverted Microscope from Zeiss. Initial processing such as contrast and brightness were completed within the AxioVision Special Edition 64-bit software. Further processing, such as removal of debris outside of the well plate and compiling of fluorescent cell images, was completed within Adobe Photoshop CC 2018. Tensile Testing Samples were fabricated in non-sterile conditions with the same flow rate and core and sheath fluid concentrations as the cell culture samples. Like in cell culturing samples, 6% alginate was diluted down to the desired 1.5% concentration with cell culture media. However, since cell viability was not being studied, water baths featured DI water instead of PBS. Fibers were dried and mounted with Epoxy glue on 15 mm wide paper frames to enable proper gripping from the tensile tester. Frames were carefully cut before testing to



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ensure that only the microfiber was tested. For this characterization, an Instron Universal Testing machine (Model 5569, Instron Engineering Corp., Canton, MA) was used with a 10 N load cell and an extension rate of 1 mm/min. A minimum of three specimens were tested from each sample type. To test wet fibers, the same procedure was utilized. Fibers dried very quickly and therefore had to be re-hydrated shortly before testing. This was accomplished with a gentle misting of DI water using a spray bottle and allowing the fibers to sit for a minimum of fifteen minutes before testing. The same tensile procedures were used. Cell-laden fibers were fabricated and mounted immediately, and were not incubated. Porosity Testing Microfibers were created in non-sterile conditions onto plain copper frames with a minimum of 5 fibers per frame. Fibers were dried for two days at room temperature. Dry weights were determined, then samples were allowed to soak within DI water and were weighed every 12 hours until their weight stopped rising, which corresponded to 24 hours. At this point, their wet weights were collected, and their porosity was calculated using Equation 1, where 𝑀𝑤 is the wet weight, 𝑀𝑑 is the dry weight, 𝜌 is the density of the liquid, and 𝑉 is the wet volume of the fiber. The volume of the microfiber samples was determined by measuring the average diameter of the fiber and approximating it as a cylinder with a height equal to the width of the copper frame, multiplied by the number of fibers on the frame. 𝑃𝑜𝑟𝑜𝑠𝑖𝑡𝑦 =

𝑀𝑤 ― 𝑀𝑑

(1)

𝜌𝑉

To generate size and shape data, samples were created in non-sterile conditions and were allowed to dry at room temperature overnight. Scanning Electron Microscopy (SEM) analysis using a JCM-6000 NeoScope Benchtop SEM with an accelerating voltage of 15 kV generated both the images and length data discussed below.


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3. Results and Discussion Synthesis of Cell-Encapsulated Alginate Fibers In the course of this study, alginate fibers were successfully fabricated with a variety of flow rates and in multiple water baths. Since the fibers were created and strong enough to gather within a water bath of pure, sterilized cell culture water, it is sure that this method relies upon ionic polymerization within the microfluidic device, unlike extrusion-based method wherein the fiber is polymerized upon reaching the collection bath. The properties of the fibers were modified by adjusting the concentration of CaCl2 · 2H2O in the water bath, thereby changing bond strength and cross-linking density, leading to a direct change in their mechanical properties. To display this, Figure 1 shows the fibers 12 hours after fabrication, and Figure 2 shows the fibers in the same wells 48 hours after encapsulation. Likewise, Figures 7 through 10 show cell-laden fibers at various time points throughout the study. Cell density within the fibers were inconsistent, as can be seen in Figures 7-10. This can be attributed to cells sinking in the syringe throughout the experiment, thereby yielding fibers with more cells per length early in the experiment when compared with fibers created towards the end. Future work can combat this inconsistency by introducing rotation to the syringe, thereby minimizing the amount of cells which are allowed to settle to the bottom of the syringe and maximizing the amount which remain in the solution and can be encapsulated within the fibers. Size Characterization of Fibers Fiber sizes and cross-sectional characteristics of dried fibers were analyzed with the use of an SEM. As seen in Figure 3, both the shape and size of the fibers were affected by the flow rate ratio and bath concentrations. This data was crucial for completing the tensile studies of dry fibers, since the cross-sectional area is a necessary piece of data for those calculations. As expected, fibers which were fabricated into a 5% CaCl2 · 2H2O bath were significantly smaller than those fabricated into a 0% CaCl2 · 2H2O bath; this is due to the increased amount of exposure to the ionic cross-linker, which causes stronger and shorter cross-linked bonds on the surface of the fibers, thereby yielding fibers with smaller crosssectional lengths and widths. Likewise, the flow rate ratio changed the length and width of the fibers for the 5% CaCl2 · 2H2O bath samples, with the 8.3 FRR (125/15 (µL/min:



µL/min) sample) showing a significant increase in size over the 10 FRR (100/10 (µL/min: µL/min)) and 12.5 (125/10 (µL/min: µL/min)) fibers, with insignificant changes for the 10 FRR and 12.5 FRR samples. Due to the fluid dynamics within the microfluidic device, the fibers are not thoroughly round, but have a longer and shorter diameter. This is reflected in Figure 3, where both ‘length’ and ‘width’ are analyzed. Porosity of Microfibers Porosity is a vital feature to consider when engineering tissue-inspired scaffolding for cell culture. High porosity allows for greater cell migration, as well as more reliable transportation of nutrients and waste into and out of the hydrogel.14 Non-porous fibers may be suitable for applications where a high degree of spatial control is desired over a tunable period of time. As seen in Figure 4, this study shows that both the concentration of CaCl2 · 2H2O in the water bath and the flow rate ratios of sheath and core fluids affected the porosity of the microfibers. This indicates that fiber porosity can be directly affected by the fabrication parameters involved with microfluidic fiber fabrication. Increasing the flow rate ratio from 8.3 (125/15 (µL/min: µL/min)) to (100/10 (µL/min: µL/min)) leads to a decrease in the porosity, just like increasing the concentration of CaCl2 · 2H2O in the bath which catches the formed fibers also decreases the porosity of the fiber. Further increasing the FRR to 12.5 (125/10(µL/min: µL/min)) did not significantly affect the porosity. The most porous fibers were created with a sheath fluid flow rate of 125 µL/min, a core fluid flow rate of 15 µL/min, and a bath concentration of 0%. Porosity of these fibers could be further increased through post-processing methods such as freeze-drying, or through pre-fabrication methods such as particle leaching.49 Elasticity Characterization of Alginate Fibers Fibers were analyzed using standard tensile testing practices, and both wet and dry samples were studied. Testing shows that dry fibers were stronger than other hydrogels used for tissue applications, including alginate-derived gels.20, 50-53 Dried fibers created into a 0% CaCl2 · 2H2O bath were weaker than those created in a 5% CaCl2 · 2H2O bath, leading to difficult manufacturing conditions which would limit their use in scale-up technologies. For this reason, while fibers created into a 0% bath were tested dry, they were omitted from the wet and wet cell-encapsulated elasticity studies.


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Results indicate that the CaCl2 · 2H2O concentration of the water bath used to collect the formed fibers significantly affects the elastic modulus of the fibers as shown in Figure 5. This shows that the degree of cross-linking of hydrogels directly influences the elastic properties of the fibers. This is due to the increased number of bonds formed in the ionic crosslinking process; for even stronger fibers, it is possible to cross-link alginate with additional crosslinking agents such as zinc (Zn2+) or other multivalent metal cations.54 For dry fibers, the flow rate ratio did not significantly affect the Young’s modulus of the fibers, likely due to the small variations in FRR chosen for this study (between 8.3 and 12.5 (µL/min/µL/min))/ For wet fibers (Figure 6), only 5% CaCl2 · 2H2O baths were used for sample preparation, due to the manufacturing difficulties of collecting fibers from pure water baths. Wet fibers were much weaker than their dry counterparts, which is confirmed in literature, and has been attributed to the plasticizing effect of water,55 which states that the glass transition temperature (Tg) is related to the weight or volume fraction (C).56 The addition of cells did not significantly affect the tensile properties of the fibers. This is likely due to the small sizes of the cells, and since the encapsulated cells were not incubated before tensile testing was conducted, cells did not have enough time to excrete extracellular matrix proteins in the volume needed to significantly affect tensile properties. Further testing over time would provide a more thorough understanding of the relationship between encapsulating cells and mechanical properties of the fibers. Literature values of the elastic properties of the in vitro human brain record Young’s modulus values of between 694 Pa57 and 9210.87 Pa,58 while the mouse brain has been reported at between 0.29-6.96 kPa.59 These values are lesser than those accomplished within this study; however, due to the highly tunable nature of the fabrication method employed, it is theoretically possible to accomplish the same in vitro elastic properties as those established experimentally. This would provide a physiologically relevant mechanical environment for cells encapsulated within the hydrogels, particularly since these factors are known to affect cell growth and differentiation.6,

60

However, for

applications where the spatio-temporal properties of the cell must be controlled and their positions manipulated, fibers with higher mechanical properties are more appropriate.



Live/Dead Cell Assays It was expected that cells deposited into a 5% CaCl2 · 2H2O bath would have a better survival rate, as pure water is known to be fatal to cells, since it causes a damaging osmotic reaction within cell walls. However, as seen in Figure 11, cells encapsulated within an alginate hydrogel fiber were able to survive complete submersion into a water bath for upwards of one minute and continue to survive for up to eleven days (Figures 10, 11). Further experiments can better elucidate cell survival under different adverse conditions. Long-term analysis showed that fibers fabricated into 5% baths with lower porosities and higher Young’s moduli maintained a higher percentage of live cells than fibers which were gathered in baths of pure DI water. This might be attributed to the difference in elastic properties of the surrounding hydrogels, which is known to affect cell health and behavior.8, 22 However, the effects of both bath concentration and flow rate ratio are minimized over time, with fibers showing insignificantly different percentages on day eleven. The mechanical properties of the fibers affected their ability to withstand in vitro conditions over time. By day 11, the fibers created at the FRR 100:10 µL/min: µL/min into a pure DI water bath were unable to withstand the live/dead cell culture protocol, and no fibers remained for analysis. Due to its fragile nature, this particular sample holds little promise for long-term applications, and performed poorly on days 1, 4 and 8 when compared with the rest of the samples. Live/dead mortality assays performed shortly after encapsulation on day one showed that the flow rate ratio 125:10 µL/min: µL/min showed a higher percentage of cell survival than the 100:10 µL/min: µL/min samples, indicating that time within the microfluidic channel is an important factor to cell health. This also implies that the alginate core fluid protects cells against the increased shear force they might experience during the encapsulation procedure.

4. Conclusions In this study, mouse astrocyte cells were successfully encapsulated within alginate fibers fabricated using a microfluidic manufacturing approach. Fibers were highly tunable by varying flow rate ratio and bath concentrations. Cells were viable within the fibers for up


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to 264 hours past manufacturing process. This proves that the fibers created in this study have potential uses for applications which require high levels of control over the spatial locations of cells, but which may need to harvest the cells again later. This also confirms the suitability of the microfluidic cell encapsulation approach due to its mild manufacturing parameters and providing appropriate fiber packing density and pore sizes. Using microfluidics, it was possible to eliminate severe manufacturing parameters such as high voltage, temperature, and surface tension associated with more common fiber fabrication techniques. This resulted in a one-day cell survival rate of 89%, with up to 60% of cells surviving within the fibers 11 days after encapsulation. Similarly, encapsulation within these alginate fibers shows promise for protecting cells from adverse conditions, such as exposure to pure water for brief amounts of time. Additionally, the mechanical properties of the fibers were readily tunable by adjusting the parameters of microfluidic fiber fabrication. Adjusting the water bath concentration lead to significant changes in the volume of pores within the fibers. For elasticity measurements, the flow rate ratio was a significant factor in deciding the Young’s Modulus, but more testing must be done to ascertain the effect of CaCl2 · 2H2O on the elastic properties of alginate fibers. As mentioned previously, porosity of fibers is an important factor to consider when designing materials for biomedical applications; in creating fibers with adjustable porosities it would be possible to have a higher degree of control over factors such as the rate of dispersion of cells, proteins or drugs from fibers into their surroundings.22, 49, 61-62 However, it would be beneficial to conduct future studies with a more robust method such as the Brunauer, Emmett and Teller (BET) technique in order to gain further insight into this parameter. This powerful technique allows for the continuous creation of fibers that have highly tunable shapes, sizes, and mechanical properties such as Young’s Moduli and porosities. In order to fully understand the conditions cells face within the hydrogel, under-utilized testing of wet cell-encapsulated fibers was used, implementing extension tests as opposed to compression testing. The factors analyzed in this study indicate that these fibers have high potential for application towards 3D cell culturing techniques. Promising results indicate that this phenomenon can be easily controlled by tuning the degree of polymerization of the outer



layers of the fiber, which might affect the ability of the cells to egress from the body of the fiber. Fibers in this study have potential uses for applications wherein cells must be precisely located and allowed to migrate from within the cell at a precise rate, such as therapeutic implantation. To accomplish this, further testing must be done in order to characterize cell migration behavior and negate the degrading effect of physiological levels of NaCl found within the body.

Acknowledgements We are grateful for critiques of this manuscript by Alex Wrede, Donald Sakaguchi, and Bhavika Patel. This work was supported by the Office of Naval Research (ONR) Grant N000141612246 and ONR Grant N000141712620. We would also like to thank Michael Cho and Rajeendra Pemathilaka for help during the project.

Conflict of Interest The authors declare no conflict of interest.

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b2

Figure 1: 1.5% alginate fibers 12 hours post-encapsulation within (a1-c1, a2-c2): 0% CaCl2 · 2H2O baths and (a3-c3, a4-c4) a 5% CaCl2 · 2H2O bath. Fibers were created with flow rates of (a1-a4): 100:10 (µL/min: µL/min); (b1-b4): 125-10 (µL/min: µL/min); and (c1-c4): 125-15 (µL/min: µL/min). Fibers were stored in maintenance media at 37˚C in 5% CO2 until they were imaged. Cells were stained with CellTracker™ CMFDA in order to achieve fluorescence. Scale bars represent 100 µm.



Figure 2: 1.5% alginate fibers 48 hours post-encapsulation within (a1-c1, a2-c2): 0% CaCl2 · 2H2O baths and (a3-c3, a4-c4) a 5% CaCl2 · 2H2O bath. Fibers were created with flow rates of (a1-a4): 100:10 (µL/min: µL/min); (b1-b4): 125-10 (µL/min: µL/min); and (c1-c4): 125-15 (µL/min: µL/min). Fibers were stored in maintenance media at 37˚C in 5% CO2 until they were imaged. Cells were stained with CellTracker™ CMFDA in order to achieve fluorescence. Scale bars represent 100 µm.


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g

Figure 3: Sizes and shapes of dried 1.5% alginate fibers created with a variety of flow rate ratios. Longitudinal images (a2-f2) show topographical features of the fibers. Fibers were created with (a-c) 0% CaCl2 · 2H2O baths and (d-f) 5% CaCl2 · 2H2O baths in water. Flow Rate Ratios were (a1-a2, d1-d2) 100:10 (µL/min: µL/min), (b1-b2, e1-e2) 125:10 (µL/min: µL/min), and (c1-c2, f1-f2) 125:15 (µL/min: µL/min). (g) Average sizes of fibers the crosssectional area for each FRR and gathering bath concentration. Cross-sectional areas of the fibers (a1-f1) are not round, and therefore have been characterized by both ‘Length’



(longer diameter) and ‘Width’ (shorter diameter) parameters. To generate the graph, n=7 samples were compared with the statistical software R, and were plotted in Origin. Values represent the mean ±1 Standard Deviation.


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Figure 4: Porosity (%) of 1.5% alginate microfibers made with a variety of Flow Rate Ratios and CaCl2 · 2H2O concentrations for the water baths. A minimum of n=5 samples were studied for this analysis. Error bars represent ±1 Standard Deviation.



a1

a2

0% CaCl2 · 2H2O Bath


b1

b2

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Figure 5: (a1-b1) Stress-strain curves of fibers created with a variety of flow rates into baths of 0 (a) and 5% (b) CaCl2 · 2H2O. (a2-b2) Average Young’s Modulus for 0% (a) and 5% (b) CaCl2 · 2H2O baths. A minimum of n=3 samples were studied for this analysis. Error bars represent ±1 Standard Deviation.


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a

b

Figure 6: (a) Stress-strain curves for wet 1.5% alginate fibers created in 5% CaCl2 · 2H2O baths with a variety of flow rates, which have units of µL/min: µL/min. Plain fibers (NC) and Cell-encapsulated fibers (C) were analyzed. Representative curves for each sample are depicted. (b) Analyzed Young’s modulus (MPa) values for the fibers shown in (a). Both standard and cell-encapsulated fibers were analyzed. Error bars represent ±1 Standard Deviation.



a1

a2

b1

b2

c1

c2

d1

d2

Figure 7: 1.5% alginate fibers 24 hours post-encapsulation with FRRS of (a1-a2, b1-b2): 100:10 (µL/min: µL/min); (c1-c2, d1-d2): 125-10 (µL/min: µL/min). Upon fabrication, fibers were gathered into a (a1-a2, c1-c2): 0% CaCl2 · 2H2O bath and (b1-b2, d1-d2) a 5% CaCl2 · 2H2O bath at room temperature. Fibers were stored in maintenance media at 37˚C in 5% CO2 until they were imaged. Live cells were stained with CellTracker™ CMFDA (Green),


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whereas dead cells were stained with Propidium Iodide (Red). Scale bars represent 100 µm.



a1

a2

b1

b2

c1

c2

d1

d2

Figure 8: 1.5% alginate fibers 96 hours post-encapsulation with FRRS of (a1-a2, b1-b2): 100:10 (µL/min: µL/min); (c1-c2, d1-d2): 125-10 (µL/min: µL/min). Upon fabrication, fibers were gathered into a (a1-a2, c1-c2): 0% CaCl2 · 2H2O bath and (b1-b2, d1-d2) a 5% CaCl2 · 2H2O bath at room temperature. Fibers were stored in maintenance media at 37˚C in 5% CO2 until they were imaged. Live cells were stained with CellTracker™ CMFDA (Green),


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whereas dead cells were stained with Propidium Iodide (Red). Scale bars represent 100 µm.



a1

a2

b1

b2

c1

c2

d1

d2

Figure 9: 1.5% alginate fibers 192 hours post-encapsulation with FRRS of (a1-a2, b1-b2): 100:10 (µL/min: µL/min); (c1-c2, d1-d2): 125-10 (µL/min: µL/min). Upon fabrication, fibers were gathered into a (a1-a2, c1-c2): 0% CaCl2 · 2H2O bath and (b1-b2, d1-d2) a 5% CaCl2 · 2H2O bath at room temperature. Fibers were stored in maintenance media at 37˚C in 5% CO2 until they were imaged. Live cells were stained with CellTracker™ CMFDA


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(Green), whereas dead cells were stained with Propidium Iodide (Red). Scale bars represent 100 µm.



a1

a2

b1

b2

c1

c2

Figure 10: 1.5% alginate fibers 264 hours post-encapsulation with FRRS of (a1-a2): 100:10 (μL/min: μL/min); (b1-b2, c1-c2): 125-10 (μL/min: μL/min). Upon fabrication, fibers were gathered into a (b1-b2): 0% CaCl2 · 2H2O bath and (a1-a2, c1-c2) a 5% CaCl2 · 2H2O bath. Fibers were stored in maintenance media at 37 ̊C in 5% CO2until they were imaged. Live cells were stained with CellTracker™ CMFDA (Green), whereas dead cells were stained with Propidium Iodide (Red). Scale bars represent 100 μm.


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Figure 11: Results of a numerical live/dead cell assay, with the percent of live cells calculated at 24, 96, 192, and 264 hours post-encapsulation. A minimum of 50 cells were assayed per fiber sample. Error bars represent ±1 Standard Deviation.



Graphical Abstract


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Microfluidic Manufacturing of Alginate Fibers with Encapsulated

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