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Published in final edited form as: Adv Mater Technol. 2023 Feb 28;8(10):2201724. doi: 10.1002/admt.202201724

Controlling the Stem Cell Environment Via Conducting Polymer Hydrogels to Enhance Therapeutic Potential

Sruthi Santhanam 1,++, Vivian R Feig 2,+, Kelly W McConnell 3,++, Shang Song 4,++, Emily E Gardner 5, Jainith J Patel 6, Dingying Shan 7, Zhenan Bao 8,9, Paul M George 10,11
PMCID: PMC13229573  NIHMSID: NIHMS2170615  PMID: 42238535

Abstract

Stem cells are a promising treatment option for various neurological diseases such as stroke, spinal cord injury, and other neurodegenerative disorders. However, the ideal environment to optimize the therapeutic potential of the cells remains poorly understood. Stem cells in the native environment are influenced by a combination of mechanical, chemical, and electrical cues for proliferation and differentiation. Because of their controllable properties, conductive hydrogels are promising biomaterials to interact with stem cells. Herein, this work develops an interpenetrating conducting polymer hydrogel with tunable mechanical properties. The hydrogel serves as a platform to provide mechanical and electrical cues for interactions with mesenchymal stem cells (MSCs). This work optimizes the formulation of the hydrogel for maximum viability of MSCs and relatively higher cytoskeletal protein expression. The viability of cells is not affected due to electrical stimulation (ES). Further, ES alters the trophic factor secretion of MSCs, with significant increase in VEGF pathway genes—VEGFA and HSPB1. In addition, substrate stiffness of the hydrogel enhances the VEGFB secretion compared to control. Hence, the conducting polymer hydrogel system creates a tunable physical and electrical niche to enhance the therapeutic potential of stem cells for neurological injuries.

Keywords: electrical stimulations, interpenetrating conducting polymer hydrogels, stem cells, tropic factor secretion, VEGF pathway

1. Introduction

In recent years, mesenchymal stem cells (MSCs) have garnered much attention for their therapeutic properties in neural environments through neurogenic differentiation and secretion of neurotrophic factors.[1,2] MSCs demonstrate some advantages over neural stem cells (NSCs) due to their low immunogenicity and immunosuppressive properties.[3] Clinical trials with MSCs have been conducted in light of the low immune response induced by the cells. While the neurorestorative applications of MSCs are promising, the harsh transplantation environment often limits their therapeutic potential.[4,5]

Biopolymers such as injectable hydrogels are used to interact with stem cells.[6] Hydrogels have been found to improve the function and activity of cells in neural environments comprised of hypoxic conditions and inflammatory responses.[7–10] Tunable hydrogel systems have been investigated as biocompatible tools for designing artificial stem cell niches. Physical factors such as matrix stiffness can alter cell properties including morphology, adhesion, proliferation, and differentiation.[11,12] Electrical stimulation (ES) also has a profound impact on modulating stem cell behavior.[13–16] Exogenous ES has shown to improve MSC viability, proliferation, and differentiation to neural lineages.[17,18] Moreover, we have demonstrated that ES of stem cells promotes the release of paracrine factors.[16,19,20]

Conducting polymer hydrogels are promising biomaterials to electrically interact with stem cells because of their controllable properties.[21,22] In particular, the electrically conductive polymer poly(3,4-ethylenedioxythiophene) (PEDOT) shows promise for neurodegenerative therapies, as PEDOT derivatives increase differentiation to neurons, neurite expression, and proliferation (Table S1, Supporting Information).[23–26] Additionally, PEDOT can be made into low-stiffness hydrogels via complexation with hydrophilic dopants, the most common of which is polystyrene sulfonate (PSS).[27–29] Compared to ionically conductive hydrogels, electrically conductive hydrogels provide a lower cell-hydrogel impedance at a broader frequency range, which provides more effective stimulation with lower voltage.[30,31] However, PEDOT-containing hydrogels often have low electrical conductivity because the hydrogel morphology and water content disrupt the PEDOT conjugated π backbone’s packing density and orientation, disrupting long-range electrical charge transport.[21,29,32] To achieve a favorable biointerface, there is a need for a low-stiffness hydrogel with high electrical conductivity. Since the required stiffness for a particular type of cell line may be different, the ability to systematically vary the stiffness while maintaining high electrical conductivity is desirable. However, the reported approaches to prepare conducting polymer hydrogels usually result in a decrease in electrical conductivity when the modulus of the hydrogel is decreased.

Given the unsolved difficulties aforementioned, a conducting polymer hydrogel with variable stiffness but maintaining stable conductivity is highly desired to maintain stem cell viability and enhance neurotrophic factor secretion by controlling the mechanical and electrical environment of stem cells. In this work, we developed a conducting polymer hydrogel with an interpenetrating network (IPN) design that allows us to systematically vary substrate mechanics while maintaining electrical conductivity. Our network comprises a PEDOT:PSS hydrogel with enhanced conductivity through acetic acid treatment. This network is interpenetrated with calcium cross-linked alginate, a nonconducting polysaccharide, to modulate the stiffness of the overall material. Because of the IPN design, we achieved a constant conductivity while tuning stiffness in the range of 1 to 10 kPa by varying concentration of the alginate. The biocompatibility of the hydrogels was determined based on the viability and proliferation of MSCs. Subsequently, the hydrogel formulation with optimal stiffness for increased viability and cytoskeletal protein expression was chosen for ES studies. Finally, the change in cell viability and neurotropic factors secreted by the cells in response to ES were determined.

2. Results and Discussion

2.1. Conductive Hydrogel Development

Because of their high water content, hydrogels possess tissue-level stiffness, making them useful substrate materials for providing stem cell support. To assess the impact of substrate mechanics and ES on stem cell behavior, conductive IPN (C-IPN) hydrogels were synthesized. The IPN design enables stiffness to be independently tuned from electrical conductivity.[27]

C-IPN hydrogels were made by interpenetrating PEDOT:PSS hydrogels with a secondary network made from calcium-cross-linked alginate (Scheme 1). Alginate is a biocompatible material commonly used in cell scaffolds.[33–37] In our preparation, varying quantities of sodium alginate were dissolved in 1.1 wt% aqueous PEDOT:PSS solution, which was then formed into a weak hydrogel through the addition of 10 vol% of 10× concentrated phosphate buffered saline (PBS).[6] The hydrogels were subsequently immersed in an acetic acid solution, which induces further phase separation of PEDOT from PSS. This enables the formation of an electrically conductive network that possesses a low frequency-independent impedance profile with impedance less than 10 Ohms in the range of 0.1–1000 Hz (Figure 1A).[6] The impedance decreased with acetic acid concentration (Figure 1B), which also increased both the storage and loss moduli of the resulting conductive hydrogel (Figure 1C, Figure S1A, Supporting Information), while their tan delta values, which quantify the ratio of loss to storage modulus, remained around 0.1 (Figure S1B, Supporting Information). The correlated increase in storage modulus and conductivity shows that acetic acid treatment results in the formation of new cross-linking interactions between conjugated PEDOT polymers, which interact with each other via π–π stacking.[27] An acetic acid concentration of 60 vol% was selected because it resulted in a PEDOT:PSS hydrogel with a storage modulus of 1 kPa, the minimum value we were interested in testing to imitate the stiffness of the neural environment.[38]

Scheme 1.

Scheme 1.

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Schematic of the preparation of poly(3,4-ethylenedioxythiophene):polystyrene sulfonate (PEDOT:PSS)/alginate conductive interpenetrating network (C-IPN) hydrogels, and chemical structures of key components (PEDOT, PSS, alginate, and Ca2+) in C-IPN hydrogels.

Figure 1.

Figure 1.

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Characterization of poly(3,4-ethylenedioxythiophene):polystyrene sulfonate (PEDOT:PSS)/alginate conductive interpenetrating network (C-IPN) hydrogels. A) PEDOT:PSS hydrogels interpenetrated with alginate possess a low frequency-independent impedance profile, maintaining the same order of magnitude of impedance despite varying concentrations of copolymer—alginate. B) Acetic acid treatment results in low, frequency-independent impedance that is approximately independent of the alginate concentration. C) Acetic acid treatment increases the storage modulus of the interpenetrating hydrogel. D) The alginate concentration can be used to significantly tune the storage modulus of the IPN hydrogels.

To form the secondary polymer network, the hydrogel next underwent immersion in aqueous solutions of calcium chloride (CaCl2). CaCl2 readily diffused into the porous PEDOT:PSS network to cross-link the alginate polymers. By varying the concentration of the stiffer alginate from 0 to 20 mg mL−1, the storage modulus of the C-IPN hydrogels was further tuned within the range of 1–10 kPa (Figure 1D, Figure S1C, Supporting Information). As alginate concentration increased, both the elastic and viscous contributions to the gels’ shear response increased, as indicated by the increase in both storage and loss moduli (Figure 1D, Figure S1D, Supporting Information). This is because incorporating more alginate into the conductive gel increases both the number of ionic cross-linking sites and the density of weak interchain interactions such as entanglements.[39] Moreover, the relative contribution of the viscous response increased with copolymer concentration, as seen by the increase in tan delta values (Figure S1E, Supporting Information).

Across this range of mechanical properties, the electrical impedance was maintained within the same order of magnitude despite the poor conductivity of alginate (Figure 1A, Figure S1F, Supporting Information), which can be attributed to the IPN design of the material. Thus, an optimum storage modulus could be selected without compromising ES efficiency.

While conductive hydrogels are typically good ionic conductors, they often possess frequency-dependent impedances with higher impedance (reduced conductivity) in the low frequency regime, reflective of poor electrical conductivity.[28] The PEDOT:PSS phase separation induced by acetic acid treatment results in an electrically conductive network bridged by conductive π-π stacking junctions between conjugated PEDOT polymer chains. This morphology enables C-IPN hydrogels to exhibit a low and frequency-independent impedance. C-IPN hydrogels therefore can support effective ES of cells across a wide range of frequencies.

2.2. Conductive Hydrogel Stiffness Alters Stem Cell Viability

To evaluate the impact of substrate stiffness on viability and growth of stem cells, we performed viability assays on MSCs seeded on the C-IPN hydrogels. Varying stiffness of the C-IPN hydrogel substrate serves as a mechanical stimulus. As assessed by a live/dead assay, all the C-IPN hydrogel groups showed comparable viabilities of MSCs on day 1. The viability of cells on C-IPN hydrogels with 15 mg mL−1 alginate (10 kPa) did not significantly change on day 7, (Figure 2A,B). AlamarBlue assays were also applied to measure the cellular viability and proliferation for over 7 days. On all C-IPN hydrogel formulations, cell densities on day 7 were 2 to 4 times higher than the initial density (day 1). Overall, the MSCs were able to proliferate across C-IPN hydrogels with varying stiffness with some indication that the 15 mg mL−1 is optimal for cellular viability.

Figure 2.

Figure 2.

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Viability and proliferation of mesenchymal stem cells (MSCs) cultured on 2D poly(3,4-ethylenedioxythiophene):polystyrene sulfonate (PEDOT:PSS)/alginate conductive interpenetrating network (C-IPN) hydrogels. A) Viability and proliferation of bone-marrow derived MSCs (BM-MSCs) cultured on 2D C-IPN hydrogels measured via live/dead assay. B) Cell count obtained from live/dead assay. The viability of MSCs were maintained through 7 days when cultured on C-IPN hydrogels with 15 mg mL−1 alginate. C) Viability and proliferation of BM-MSCs cultured on 2D C-IPN hydrogels measured via alamarBlue assay. Control is cells alone cultured on a tissue culture plastic substrate. Experimental groups are cells cultured on the surface of C-IPN hydrogels of varying concentrations of alginate. Analyzed using a two-way ANOVA, followed by Tukey’s HSD post hoc test with *p < 0.05. N = 3, Data represented as mean ± S.D.

Substrate stiffness affects the cytoskeletal structure of the cell.[40–42] To determine how the change in substrate stiffness affects the MSCs, we stained for cytoskeletal proteins, F-actin and RhoA. Immunocytofluorescence images reveal that the MSCs had a circular morphology on the C-IPN hydrogel substrate (Figure 3A). Further, substrate stiffness changes altered the fluorescence of RhoA and F-actin in the cell (Figure 3B). The RhoA in an MSC was found to be significantly higher in C-IPN hydrogel of 15 mg mL−1 compared to other hydrogels. F-actin also followed a similar trend as RhoA, however without any significant difference between the groups. RhoA activation is thought to interact with either YAP1 (YAP signaling) or ROCK (RhoA/ROCK signaling) that are vital for modulating cellular behavior and fate.[40,43,44] Taken together, a small change in substrate stiffness can alter the proliferation and regulation of RhoA. Therefore, this suggests that the ability to fine tune hydrogel stiffness is important to modify cell properties.

Figure 3.

Figure 3.

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Immunofluorescence staining for cytoskeletal proteins. A) Immunofluorescent images illustrating RhoA and F-actin expression of mesenchymal stem cells (MSCs) via poly(3,4-ethylenedioxythiophene):polystyrene sulfonate (PEDOT:PSS)/alginate conductive interpenetrating network (C-IPN) hydrogels (alginate 0–20 mg mL−1). B) Quantification for fluorescent area per cell determined using imageJ. Analyzed using one-way ANOVA followed by Turkey’s post hoc test for multiple comparisons with *p < 0.05. N = 3, Data represented as mean ± S.D.

2.3. Electrical Stimulation Modulates Stem Cell’s Trophic Factor Secretion

ES changes the transcriptome of the cells and enhances neurotrophic factor release in neural stem cells.[1,2] All the ES experiments were performed on the hydrogel that maintained better cellular viability and proliferation—PEDOT:PSS/alginate (15 mg mL−1 alginate). To make sure that ES does not affect cell viability, we performed live/dead assays before and after stimulation. The MSCs were seeded on top surface of the C-IPN hydrogel (15 mg mL−1 alginate) and were electrically stimulated using a custom-made plate–plate setup for 1 h at 20 Hz and ±1 V alternating current (AC) (Figure 4A). We optimized ES parameters based on our lab’s previous work and the requirement for minimal frequency that produces stable conductivity through the hydrogel but not PBS (ionic conductivity). The amplitude and frequency across the hydrogel was also measured simultaneously to verify the stimulation (Figure S2, Supporting Information). The live/dead assay before and after 1 h of ES revealed that the MSCs were viable after ES via the C-INP hydrogel (Figure 4B). When quantified, there were no significant difference in the percentage viability between the groups (p value = 0.7279) when analyzed using a nested t-test, thereby confirming that ES does not affect viability of the cells.

Figure 4.

Figure 4.

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Electrical stimulation (ES) and cell viability. A) Schematic of custom-made plate–plate type ES setup. The cells cultured on the top surface of the conductive interpenetrating network (C-IPN) hydrogel were electrically stimulated (ES) using the plate–plate type conductive electrodes for 1 h on day 1. B) Viability of mesenchymal stem cells (MSCs) before and after stimulation via live/dead assay. C) Percentage viability quantified based on live/dead assay. There was no significant difference between before and after stimulation when analyzed using a nested t-test. N = 3, and data represented as mean ± S.D.

To understand the transcriptome changes to the MSCs due to ES, the relative gene expression levels for various paracrine factors with respect to the housekeeping gene were determined using qRT-PCR (Figure 5, Figure S3, Supporting Information). The experimental groups include electrically stimulated MSCs on the C-IPN hydrogel (15 mg mL−1 alginate, ES), unstimulated MSCs on the C-IPN hydrogel (US), and control (cells on a tissue culture plastic (TSP) substrate). B2M, a commonly used housekeeping gene, was chosen as the housekeeping gene for these experiments since it maintained nearly constant Ct values across the experimental parameters. Determining the relative gene expression changes via ΔΔCt method revealed that the electrically stimulated MSCs had significantly higher (nearly fourfold) gene expression of VEGFA and HSPB1 genes compared to controls (Figure 5A,B). HSPB1 is a heat shock protein that includes the HSP27 gene. HSP27 is a downstream protein in the VEGFA pathway that is critical for actin reorganization and angiogenesis. On the other hand, nearly threefold and over threefold increase in VEGFB gene was observed in MSCs of unstimulated and electrically stimulated group, respectively, compared to that of TSP control (Figure 5C). Higher expression of VEGF-B in the US hydrogel compared to the TSP control is indicative of the effect of substrate stiffness from the C-IPN hydrogel. Alternatively, FGF2 and other paracrine related genes were not changed by the hydrogel or ES (Figure 5D, Figure S3, Supporting Information). Taken together, the VEGF pathway in the MSCs has been significantly modulated due to electrical and mechanical cues from the conductive hydrogel.

Figure 5.

Figure 5.

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Gene expression changes due to electrical stimulation (ES) of mesenchymal stem cells (MSCs) via conductive interpenetrating network (C-IPN) hydrogel discs. A–D) Relative gene expression of target genes VEGFA, VEGFB, FGF2, and HSPB1, respectively, with respect to housekeeping gene—B2M when measured using quantitative real-time polymerase chain reaction (qRT-PCR). Analyzed using a one-way ANOVA, followed by Tukey’s HSD post hoc test with *p < 0.05, **p < 0.01, and ****p < 0.0001. N = 3–6, Data represented as mean ± S.D.

VEGF signaling has thought to play a critical role in various neurological injury and neurodegenerative diseases such as spinal cord injury, traumatic brain injury, stroke, epilepsy, and Parkinson disease.[11,16,45] VEGF has trophic effects in neurological diseases mainly cause of the intertwined interaction (or cross-talk) between angiogenesis and neurogenesis.[46] Previously, our lab has shown that ES of neural stem cells upregulates the VEGF secretion, which subsequently enhances the recovery after stroke in rodents.[16] Similarly, VEGF gene delivery via nonviral carriers in rat models of Parkinson’s disease (PD) prevented loss of motor function due to PD, and loss of dopaminergic neurons and nerve fibers.[47] MSCs have been used as a stem cell therapy for various neurodegenerative diseases and delivery of electrically stimulated MSCs with enhanced VEGF secretion is a promising treatment option.

3. Conclusions

A conductive C-IPN hydrogel disc, developed by interpenetrating PEDOT:PSS and alginate polymers, serves as a platform to modulate the mechanical and electrical stimuli of stem cells. The hydrogels demonstrated nearly constant impedance independent of changes in frequency and stiffness. The MSCs were found to be viable and proliferating in substrates of optimized modulus. In addition, MSCs survived well after ES via C-IPN hydrogel. The VEGF pathway genes was found to be augmented due to mechanical (VEGF-B) and electrical cues (VEGF-A, HSPB1). Hence, the conductive hydrogel system creates a tunable physical and electrical niche to enhance the therapeutic potential of stem cells for its use in neurological injuries.

4. Experimental Section

Materials:

PEDOT:PSS (Orgacon ICP 1050) was provided as a 1.1% dispersion in water by Agfa-Gevaert N.V. and was filtered through a 0.2 micron Nylon filter prior to use. Alginic acid sodium salt (CAS #9005–38-3), acetic acid (CAS #64–19-7), and calcium chloride (CAS #10043–52-4) were purchased from Sigma–Aldrich. Human bone-marrow derived mesenchymal stromal cells (BM-MSC, catalog #70071) were purchased from STEMCELL Technologies (Vancouver, Canada). All cell culture reagents were obtained from ThermoFisher Scientific (Waltham, MA) or STEMCELL Technologies. The cell culture plates and flasks were purchased from Genesee Scientific (San Diego, CA).

Synthesis of Conducting Hydrogels:

In a typical synthesis for conductive IPN hydrogels, a desired quantity of sodium alginate was first dissolved in the PEDOT:PSS dispersion. Next, 90 vol% of this mixture was combined with 10 vol% of 10× concentrated pH 7.4 PBS solution while stirring vigorously with a magnetic stir bar. Alginate concentrations refer to the quantity of sodium alginate relative to the final volume of the mixture (including PBS). After removing the stir bar, the resulting hydrogel was submerged in an equivalent volume of acetic acid aqueous solution for 24 h to form the electrically conductive hydrogel network. Finally, the alginate network was cross-linked by immersing the hydrogel in a solution of 11 mg mL−1 CaCl2 in 1× PBS for 24 h, exchanging the solution multiple times until a neutral pH was obtained. Hydrogels were stored in the CaCl2-containing PBS mixture until ready to use.

Material Characterizations:

Oscillatory rheology was performed using a TA Instruments ARES-G2 rheometer using 8 mm parallel plates at 37 °C. Electrochemical impedance spectroscopy (EIS) measurements were performed using a Bio-Logic VSP potentiostat by applying an AC voltage of amplitude 10 mV. Hydrogel samples (diameter = 0.4375 inch, thickness = 0.125 inch) for EIS were prepared using a tissue punch and sandwiched between gold-plated glass slides.

Cell-Laden Hydrogel Development:

BM derived mesenchymal stromal cells (BM-MSC), were cultured and passaged as per the protocol from STEMCELL Technologies. The MSCs (P4-P6, about 30 000 cells per hydrogel) were seeded on the top surface of the hydrogel (N = 3 for all experiment groups). The cell-hydrogel samples were left undisturbed at room temperature for 30 min and then maintained in cell culture media (89% v/v DMEM/F:12, 10% v/v fetal bovine serum, 1% v/v penicillin/streptomycin) at 37 °C for viability and immunocytochemistry. The experimental groups were PEDOT/alginate hydrogels made from a 1.1 wt% PEDOT:PSS dispersion and varying concentrations of alginate—10, 15, and 20 mg mL−1. Controls were PEDOT:PSS hydrogel alone without any alginate (0 mg mL−1) and cells alone on a coverslip.

Live/Dead Assay:

On day 2, the samples were rinsed with 1× PBS, and incubated with ethidium homodimer-1 and calcein AM (2 μL mL−1 of 1× PBS) for 20 min at 37 °C. After incubation, the cells were rinsed with 1× PBS, and imaged at constant exposure levels using a fluorescent microscope (Keyence BZ-X700, Itasca, IL). Three images were obtained at different parts of the hydrogel. The viability was quantified by measuring the total area under green fluorescence for a section of the image after thresholding for a constant background using ImageJ.

AlamarBlue Assay:

The protocol of alamarBlue assay for cell proliferation on cell-laden hydrogel samples was developed and modified based on a protocol from Thermofisher Scientific and two reference articles.[33,34] The cell-laden hydrogel samples were immersed in a 10% alamarBlue cell viability reagent and incubated at 37 °C for 3, 24, and 48 h in the dark. At different time periods, 20 μL per sample was removed and mixed with 80 μL of 1× PBS in a 96-well plate. The duplicates of each sample were measured for absorbance at 570 and 600 nm using a multiplate reader (SpectraMax, Molecular Devices, CA). The percentage reduction in absorbance (percentage viability) was calculated with respect to the initial measurement (3 h incubation) as per the manufacture protocol.

Immunocytochemistry:

The MSCs seeded on the hydrogel were stained for cyto-skeletal proteins such as F-actin and Rho-A that change with substrate stiffness. Briefly, the cell-laden hydrogels were immersed in 4% paraformaldehyde for 20 min at room temperature to fix the cells. The cells were then permeabilized with 0.1% Triton X-100 and 1% BSA, blocked with 5% normal goat serum (NGS), and incubated in the primary antibodies solution—F-actin and Rho-A (1:500, Thermofisher Scientific) overnight at 4 °C. The next day, the cells were washed with 1× PBS and incubated in secondary antibodies solution—Alexafluor 488 and 555 at 1:200 dilution for 2 h in dark. The cells were finally stained with DAPI (1:2000) for 5 min and stored in 1× PBS for imaging. The cell-laden hydrogels were imaged at constant exposure levels using Keyence BX700 fluorescent microscope and quantified using ImageJ after constant thresholding.

Electrical Stimulation:

For ES experiments, about 275k MSCs (P7-P10) were seeded per PEDOT:PSS/Alginate hydrogel (15 mg mL−1) as described in the cell-laden hydrogel section (N = 3 for all experiment groups). The cell-laden hydrogels were electrically stimulated using a custom-made plate–plate type system. The cells were stimulated at ± 1 V AC at 20 Hz frequency for about 1 h. The viability of cells before and after stimulation were evaluated using live/dead assay. Total RNA of the cells from the electrically stimulated hydrogel were extracted for quantitative real-time polymerase chain reaction (qRT-PCR). The control were cells on hydrogel without ES and cells alone plated in a 6-well plate.

RNA Extraction and Quantitative Real-Time Polymerase Chain Reaction Analysis:

Total RNA was extracted from the cells seeded on the top surface of the hydrogel using a Qiagen RNeasy Plus Micro Kit (Qiagen, Germantown, MD) and the first-strand cDNA was synthesized using iScript cDNA Synthesis Kit (Bio-Rad, Hercules, CA). qRT-PCR was performed to quantify gene expression using the TaqMan-polymerase and primers (Qiagen). Pre-developed TaqMan reagents were used for human VEGFA (Hs00900055_m1), BDNF (Hs02718934_s1), NRN1 (Hs02786624_g1), VEGFB (Hs00173634_m1), GDNF (Hs01931883_s1), FGF-2 (Hs00266645_m1), HSPB1 (Hs00356629_g1) and housekeeping genes were GAPDH (Hs02786624_g1), B2M (Hs00187842_m1), and ACTB (Hs03023943_g1). Housekeeping gene B2M was used as a reference gene. qRT-PCR was carried out on a QuantStudio 6 Flex Real-Time PCR System (ThermoFisher Scientific). The Delta–Delta CT method was used to determine the relative gene expression levels with GAPDH as a housekeeping gene and cells alone in a 6-well plate as reference.

Statistical Analysis:

All data are represented as the mean ± standard deviation (S.D.). N values indicate the number of independent samples. An analysis of variance (ANOVA) test was used for multicomponent comparisons (n > 3 independent variables) after the normal distribution was confirmed. Tukey post hoc analysis was performed to investigate the differences between variables.

Supplementary Material

Supplemental

Supporting Information is available from the Wiley Online Library or from the author.

Acknowledgements

The work was supported in part by the NIH grant K08NS089976 to P.M.G.; by a Big Ideas in Neuroscience grant from the Wu Tsai Neurosciences Institute at Stanford University to Z.B.; by the Department of Defense (DoD) through the National Defense Science and Engineering Graduate (NDSEG) Fellowship Program to V.R.F. Part of this work was performed at the Stanford Nano Shared Facilities (SNSF), supported by the National Science Foundation under award ECCS-1542152. We thank Dr. Kati Andreasson (Stanford University) for use of the qRT-PCR.

Footnotes

Conflict of Interest

The authors declare no conflict of interest.

Contributor Information

Sruthi Santhanam, Department of Neurology and Neurological Sciences, Stanford University School of Medicine, Stanford, CA 94305, USA.

Vivian R. Feig, Department of Materials Science and Engineering, Stanford University, Stanford, CA 94305, USA.

Kelly W. McConnell, Department of Neurology and Neurological Sciences, Stanford University School of Medicine, Stanford, CA 94305, USA.

Shang Song, Department of Neurology and Neurological Sciences, Stanford University School of Medicine, Stanford, CA 94305, USA.

Emily E. Gardner, Department of Neurology and Neurological Sciences, Stanford University School of Medicine, Stanford, CA 94305, USA

Jainith J. Patel, Department of Neurology and Neurological Sciences, Stanford University School of Medicine, Stanford, CA 94305, USA

Dingying Shan, Department of Neurology and Neurological Sciences, Stanford University School of Medicine, Stanford, CA 94305, USA.

Zhenan Bao, Department of Materials Science and Engineering, Stanford University, Stanford, CA 94305, USA; Department of Chemical Engineering, Stanford University, Stanford, CA 94305, USA.

Paul M. George, Department of Neurology and Neurological Sciences, Stanford University School of Medicine, Stanford, CA 94305, USA Stanford Stroke Center, Stanford University School of Medicine, CA, USA.

Data Availability Statement

The data that support the findings of this study are available from the corresponding author upon reasonable request.

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Associated Data

This section collects any data citations, data availability statements, or supplementary materials included in this article.

Supplementary Materials

Supplemental
NIHMS2170615-supplement-Supplemental.pdf (1.1MB, pdf)

Data Availability Statement

The data that support the findings of this study are available from the corresponding author upon reasonable request.

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