Wirelessly Powered-Electrically Conductive Polymer System for Stem Cell Enhanced Stroke Recovery

Abstract

Effective stroke recovery therapeutics remain limited. Stem cell therapies have yielded promising results, but the harsh ischemic environment of the post-stroke brain reduces their therapeutic potential. Previously, we developed a conductive polymer scaffold system that enabled stem cell delivery with simultaneous electrical modulation of the cells and surrounding neural environment. This wired polymer scaffold proved efficacious in optimizing ideal conditions for stem cell mediated motor improvements in a rodent model of stroke. To further enable preclinical studies and enhance translational potential, we identified a method to improve this system by eliminating its dependence upon a tethered power source. We have herein developed a wirelessly powered, electrically conductive polymer system that eases therapeutic application and enables full mobility. As a proof of concept, we demonstrate that the wirelessly powered scaffold is able to stimulate neural stem cells in vitro, as well as in vivo in a rodent model of stroke. This system modulates the stroke microenvironment and increases the production of endogenous stem cells. In summation, this novel, wirelessly powered conductive scaffold can serve as a mobile platform for a wide variety of therapeutics involving electrical stimulation.

Graphical Abstract

A wirelessly powered conductive polymer scaffold has been developed to enhance stroke recovery using stem cells. As a proof-of-concept, the scaffold enabled stem cell delivery and stimulation in a rodent model of stroke, and increased endogenous stem cell production. This mobile platform eliminates the need for a tethered power source, facilitating wider therapeutic applications involving electrical stimulation.

1. Introduction

Electrical stimulation of stem cells has emerged as a promising technique for cellular modulation. This approach is based upon the intrinsic electrophysiological nature of most stem cells, which respond to electrical cues in their environment (1). The response to these electrical signals can manifest in various ways, ranging from altered cell morphology and migration, to changes in cellular fate, function, and microenvironment (2–4).

Conductive polymers have come to the forefront as vital tools for efficient electrical modulation of stem cells due to their unique properties (5). Polymers like polypyrrole (PPy), polyaniline (PANI), and poly(3,4-ethylenedioxythiophene) (PEDOT) exhibit excellent electrical conductivity, flexibility, and biocompatibility, thus making them ideal for biological interfaces (6–8). PPy, in particular, stands out due to its ease of synthesis, exceptional conductivity, and high biocompatibility. These attributes have led to its use in promoting adhesion and differentiation of stem cells (9). While certain conductive polymers have advantages in particular aspects, such as mechanical properties and stability, the relative simplicity in synthesis, effectiveness, and biocompatibility of PPy makes it an ideal choice for many applications in stem cell electrical modulation (10).

To capitalize on these findings, our team previously developed a conductive polymer system, utilizing PPy, for the concurrent delivery of stem cells and electrical stimulation in vivo (11). Our system demonstrated that exogenous stem cells, when electrically stimulated after transplantation, can augment multiple regenerative processes, such as enhancing the proliferation of endogenous stem cells. In particular, we demonstrated that this system could significantly promote functional recovery in a rodent model of stroke (12), thereby showcasing the immense potential of these findings.

One limitation of our initial conductive polymer system and other neural modulation platforms is that they rely on tethered wiring systems for electrical stimulation. This approach often hinders mobility and ease of function, a common limitation of many therapeutic devices (13). Even systems geared for clinical use, such as those used in deep brain stimulation (DBS), are predominantly either tethered or depend upon bulky battery packs (14–16). Wireless, battery-free stimulation devices usually lack powering range, thereby immobilizing subjects (17). Recognizing these challenges, we endeavored to enhance the translational capacity of our prototype and the wearability of other tethered stimulation devices by creating a battery-free, compact wireless system.

We have herein developed a wirelessly powered, conductive scaffold that enables powering across a large arena without bulky batteries or recharging systems (Figure 1). This novel design can confer therapeutic effects via electrical stimulation of stem cells without hindering mobility. As a proof-of-concept, we demonstrate its ability to electrically stimulate human neural stem cells wirelessly in a rodent model of stroke.

Figure 1. Wirelessly powered conductive polymer system.

Experimental pipeline for the preparation of a wirelessly powered conductive scaffold system for stem cell engraftment and electrical stimulation in vivo. In this proof-of-concept design, human induced pluripotent stem cells (hiPSCs) were differentiated into neural progenitor cells (hNPCs) prior to being engrafted onto the conductive polymer scaffold. The scaffold was then implanted adjacent to the stroked brain region (peri-infarct) of rodents that underwent an induced distal middle cerebral artery occlusion (dMCAO) stroke. The novel wireless stimulator enables electrical stimulation of the stem cells on the polymer scaffold in vivo via an electromagnetic field from the environment, which allows for treatment in awake, freely moving subjects.

Our results indicate that this wirelessly powered conductive scaffold can serve as an effective platform to electrically modulate stem cells and the brain in vivo for a variety of diseases. Overall, this wireless system stands as an advancement in the field of regenerative medicine, providing a potential new therapeutic pathway for stroke and other neurological disorders with unmet clinical needs (18).

2. Results

2.1. Wireless power delivery system

A major challenge in designing wireless power delivery systems is the tradeoff between coverage and tracking complexity. Here, we utilize the metallic cage and use a radio frequency (RF) source with wavelength much larger than the dimension of mesh sizes of the cage (Figure 2a) to confine the electromagnetic field. With an appropriate placement of the feed location for the RF source (x = 15 cm, y = 0, z = 12 cm), we achieve an even power distribution (electric field strength fluctuation of approximately ±3 dB) within an arena of 30 cm × 21 cm × 21 cm without the need of any tracking mechanism (Figure 2b). The subwavelength confinement of the RF source by the cage accomplishes a minimum reflection coefficient of −9 dB, translating to only 12.5% wasted power (Figure 2c). When 1 W is fed into the RF source, the simulated average specific absorption rate (SAR) is below the limit of 1.6 W/kg (Figure 2d), confirming the safety of the system (19,20). The arena (wireless stimulation area) achieves one of the largest uniform power delivery platforms for rats.

Figure 2. Electromagnetic simulation of wireless power transmitter.

(a) Depiction of arena setup for use with a finite element method (FEM) simulator (Ansys HFSS) with defined boundary conditions, antenna port, and stimulus. (b) Results of input reflection coefficient. (c) Electrical field distribution of the arena. (d) Specific absorption rate mapped on the rat body (SAR) and (e) vector distribution of the cage transmitter at 1.3 GHz.

To achieve efficient wireless power transmission to a rodent, designing the cage with a resonance frequency that closely matches the rodent’s inherent resonance frequency is crucial. This principle resembles the phenomenon of maximum transmission efficiency in power transmission between antennas, where resonant frequencies of the antennas must be equivalent. Ho et al. (21) demonstrated that rodents have an intrinsic resonant frequency at approximately 1.3 GHz. Moreover, as the rodent has a high dielectric, its −10-dB bandwidth can extend up to 20%, allowing it to absorb high energy within the frequency range of 1 to 1.5 GHz and supply the implant. To wirelessly power the implanted device in the rodent, a commercially available cage was converted into a wireless power transfer system by confining electromagnetic energy within it. Through electromagnetic field simulation, we confirmed that the resonance frequency of the 30 cm × 21 cm × 21 cm cage was around 1.3 GHz, which is optimal for efficient power transmission to the rodent. The simulation (Fig. 2e) further revealed that the TE011 mode provides the highest power and uniform energy to the area where the rodent is present, with a grid spacing of 3 cm, which is λ/13 and small enough for the cutoff frequency, resulting in most of the energy conservation inside the cage.

2.2. Wireless stimulator development

The wireless implantable stimulator, which produces power from the evenly distributed electromagnetic (EM) field in the arena, was designed to generate biphasic stimulation waveforms. Unlike unipolar stimulation waveform, the generation of biphasic waveform requires control logics to switch the phases at the right timing. In a wired implementation, this is easy to implement; however, in a wireless system, this might require the implementation of a radio telemetry which is overly complex for a fully implantable system. Here, we propose a simplified circuitry using a level shifter implemented by a capacitor divider and an oscillator to generate biphasic waveform without the need of any complex control logic and radio telemetry. Figure 3(a) shows the circuit schematic of the device, in which, a cathode was connected to a capacitive divider with a voltage level of 0.5V_dd, whereas the anode was connected to the signal output switching between 0 and V_dd. Therefore, a stimulation pattern switching between −0.5V_dd and 0.5V_dd is applied to the polymer scaffold. A 3D model of the device is shown in Figure 3(b), and Figure 3(c) shows the measurement results of the overall efficiency from the RF source to the implantable device. An average efficiency of −21 dB (0.8 %) is achieved in a 14 cm × 14 cm area around the center of the arena and a minimum efficiency of −30 dB (0.1 %) is achieved at the edge of the arena, translating to ≥ 1mW received power from an 1-W RF source. In this particular preclinical study prototype, the arena was designed of sufficient size to encompass 4 rats at the same time.

Figure 3. Wireless stimulator design.

(a) Circuit schematic of the implantable stimulator device. (b) 3D model of the device (scale). (c) Overall power transfer efficiency from the arena to the stimulator.

2.3. In vitro cellular characterizations

To evaluate the efficacy of the wirelessly powered conductive scaffold, we tested the system in vitro (Figure 4). The wireless stimulator was connected to the electrically conductive polypyrrole (PPy) scaffold and found to have an impedance of (13.63-j13.62) KOhm at 100Hz; with a reference electrode impedance of (152-j335) Ohm at 100Hz. We cultured the PPy with human iPSC-derived neural progenitor cells (hNPCs) and stimulated (Figure 4a) with parameters that were derived experimentally from prior studies (4,8). To assess the electrically stimulated cells for viability, a live/dead assay was conducted 48h after stimulation. Results suggest that the cells were viable after electrical stimulation (Figure 4b.); which corroborates with prior study results utilizing identical parameters delivered via a wired system (8,12). Evaluation of gene expression changes after stimulation via qRT-PCR analysis revealed that certain genes implicated in stroke recovery, vascular endothelial growth factor B (VEGFB) and neuritin-1 (NRN1), which are crucial for blood vessel formation and neural regeneration (22,23), respectively, were upregulated in comparison to unstimulated control cells (Figure 4c). This data supports findings from prior work (8, 11), where a similar effect is demonstrated with the tethered stem cell scaffold system, indicating that this novel wireless powered system can confer analogous results.

Figure 4. Gene expression changes of stem cells after wireless electrical stimulation.

(a) Electrical stimulation readout via in-vitro cell chamber system measured using function generator. (b) Image of cells on polymer scaffold after electrical stimulation with live/dead staining. Green indicates live cells and red indicates dead cells. (c) Relative fold changes in gene expression when measured using qRT-PCR with GAPDH and cells alone control as reference. Analyzed using an independent t-test with a 95% confidence interval (CI) for the mean difference; n=3 per group, data represented as mean ± S.D. *p<0.05, **p<0.01.

2.4. In vivo characterization of the wireless stimulation system

To assess the feasibility of the wirelessly powered conductive scaffold in vivo, we implanted the device over the post-ischemic brain tissue of rats that underwent an experimental distal middle cerebral artery stroke (dMCAO) (Figure 5a, b). The experimental groups were those with scaffolds cultured with hNPCs, with and without electrical stimulation. The scaffold alone (no stem cells), with and without electrical stimulation served as controls. Each animal was stimulated for three days at 1 hour per day (Figure 5a); these parameters were optimized experimentally from prior studies (4,12). The rats exhibited normal behavior and were freely moving throughout the electrical stimulation period without any visible discomfort (Figure 5c, Supplemental Video 1).

Figure 5. In-vivo implantation of wireless stimulator with conductive scaffold.

(a) Experimental timeline for procedures with stroke surgery (dMCAO) as day 0, implantation of the conductive scaffold 7 days post-stroke, followed by 3 consecutive days of wireless stimulation. (b) Representative image of implanted wireless electrical stimulator and polymer scaffold with/without cells. (c) Representative image of a rat while being electrically stimulated inside the cage.

After 21 days post stroke, animals were sacrificed for cerebral infarct evaluation via cresyl violet staining (24). Histological assessment of brain slices in each experimental group revealed that all had an average infarct size in the range of 20 – 30% (Supplemental Figure 1), with no significant difference across groups.

Electrical stimulation of human neural progenitor cells has been shown to improve functional recovery after stroke and to augment production of endogenous stem cell production (12). To investigate if the same mechanism is relevant when stimulated via the wirelessly powered scaffold, we assessed the production of endogenous stem cells with neuronal fate via immunohistochemical analyses of neuroblasts (Figure 6A). To mark new cellular turnover, we injected 5-bromo-2’-deoxyuridine (BrdU), a fluorescent thymidine analog that incorporates into proliferating cells (25), intraperitoneally to the rats prior to electrical stimulation (day 7). To identify the production of proliferative neuroblasts after electrical stimulation, we assessed post-mortem brain tissue for doublecortin (DCX+) cells that co-stained positive for bromodeoxyuridine (BrdU). Neuroblasts are migratory cells, and in rodents have been reported to originate in the sub-ventricular zone (SVZ) and migrate toward the injured peri-infarct zone (PI) post-stroke (26,27). We therefore evaluated the number of DCX+BrdU+ cells in the SVZ and PI regions (Figure 6b). At 3 weeks post-stroke, the number of DCX+BrdU+ cells were statistically significantly higher in the SVZ of the wirelessly stimulated stem cell group, with nearly 2 to 4-fold higher total proliferative neuroblasts in comparison to other group (Figure 6c). There was no significant increase in endogenous stem cells in the PI region of any experimental group, however, indicating that the cells had not yet migrated toward the infarct site at 3 weeks post-stroke.

Figure 6. Wireless electrical stimulation enhanced endogenous stem cell proliferation.

(a) Representative fluorescent images of brain tissue stained for DCX (pink) and BrdU (green) in the SVZ and PI region. (b) Representative image of brain tissue with areas demarcated as SVZ and PI region highlighted red, and marked “1” and “2”, respectively. (c) Quantification of cells co-stained positive for both DCX and BrdU. Analyzed using one-way ANOVA followed by Tukey’s post hoc test; n=4 per group, data represented as mean ± S.D. *p<0.05, **p<0.01

To delineate between transplanted stem cells on the scaffold from endogenous stem cells produced intrinsically, we quantitated human nuclear antigen (HuNu) positivity via immunohistochemical analyses (Supplemental Figure 2), given that the stem cells engrafted were human derived iPSCs. The number of positive cells were negligible (average range of 2 – 4 cells per slice) in all experimental groups. This data supports findings from our prior studies, which utilized triple labelling (HuNu-DCX+BrdU+) to confirm that these transplanted stem cells minimally integrate with their host environment despite their potential to enhance regenerative processes (8,12). Taken together, this data suggests that electrical stimulation via a wirelessly powered system can increase the proliferation of endogenous neuroblasts. An increase number of neuroblasts has been shown to correlate with improved functional deficits post-stroke (12).

3. Conclusion

Here we introduce a fully implantable wirelessly powered conductive scaffold that can simultaneously modulate stem cells and their local transplantation environment, via electrical stimulation, without hindering the mobility of treatment subjects. The stimulation arena has one of the largest areas for freely behaving rats without a bulky head stage and any complex tracking mechanism. The system consists of an electromagnetic based power delivery method combined with a stimulation of a conductive polymer implant that can wirelessly harvest power and generate biphasic electrical waveforms. This electrical stimulation, in turn, has the capability to enhance both delivery and modulation of stem cells in a variety of disease models. Our findings here in particular demonstrate that stem cells transplanted with this wireless system have the potential to enhance the post-stroke neural environment and create a more favorable environment for the proliferation, migration, and differentiation of endogenous neural stem cells in the SVZ, promoting neuronal regeneration in areas affected by injury or stroke. This is achieved in freely moving rats, without the need for anesthetics or constraints.

Over recent years, developments in wearable electronics have led to the realization of FDA-approved devices for the ambulatory treatment of many diseases, from glioblastoma (28) to paralysis (29). Although further preclinical studies are needed including aged and female animals, and animals with alternative forms of immunosuppression, as well as those with common stroke co-morbidities such as hypertension and diabetes, we believe that the wireless conductive scaffold presented herein has vast potential for the enhancement of a variety of stem cell based therapeutics. Beyond improvements in stem cell delivery, for instance, it can further enable combinatorial therapeutic approaches that require mobility (30), such as with electrical stimulation of transplanted stem cells in conjunction with rehabilitative exercises for post-stroke patients. Electrical stimulation of stem cells simultaneous to physical tasks that heighten neuronal activity could potentially impact a variety of diseases with motor disorders. Future studies will aim to elucidate the therapeutic possibility of such approaches.

4. Methods

4.1. Electrically conductive scaffold system fabrication

PPy (Sigma-Aldrich) was electroplated onto indium tin oxide (ITO) slides (Delta Technologies) as described previously (8,31). Briefly, PPy was electrodeposited on an ITO plate template via an applied current density of 1mA/cm² in a 0.2M PPy and 0.2M NaDBS (Sigma-Aldrich) for 4 hours at 25°C. Platinum mesh was the reference electrode. The PPy scaffold was then removed from the ITO slide. The PPy scaffold was clamped between pieces of polydimethylsiloxane (PDMS; Sylgard, Dow) with a chamber slide forming cell chambers (Lab-Tek, Thermo Fisher) as previously described (11). Wires were attached to the PPy scaffold outside of the media containing chambers. The wireless electrical stimulator was attached to one-side of the PPy scaffold. In in vitro studies, electrical stimulation was applied one day after cell seeding at 1.6 GHz for 1 hr. In in vivo studies, cell chambers and PDMS were unclamped and separated from the PPy scaffold and wireless chip. The dimensions of the implanted PPy scaffolds were 1 × 3 × 0.25 mm³.

4.2. Wireless powering system

Electromagnetic (EM) energy in the range of GHz frequency is selected as the method for wireless powering due to the benefit of miniaturized receiver antenna and comparable high energy delivery efficiency. To accommodate free movement of test subjects during the experiment, we transformed a common rodent housing unit into a microwave resonator, where the EM energy is approximately evenly distributed in a 10-cm radius circular region with a power transfer efficiency of −21dB. Ansys HFSS was used for FEM simulation of the wireless stimulation system. The HFSS built-in local SAR calculator was used to obtain the SAR plot of the rat.

4.3. Implantable stimulator

Based on off-the-shelf electronic components, the implantable stimulator receives the EM energy through a loop antenna, generates a DC supply voltage with a charge-pump (CP) rectifier, and outputs the stimulation waveform by an oscillator and driver (inverter). The cathode is connected to a capacitive divider with a voltage of 0.5V_dd, whereas the anode is connected to the driver switching between 0 and V_dd. As a result, the equivalent stimulation waveform applying to the polymer scaffold system is biphasic switching between −0.5V_dd and +0.5V_dd. For impedance measurements, two PPY electrodes were submerged in PBS solution with distance of 5mm. The end of each electrode was connected to external wire through conductive silver epoxy and proper insulation. With small-amplitude sinusoidal stimulus of 100Hz to the external wires of the two electrodes, the serial impedance of the two PPY electrodes, Zserial, can be measured. The PPY electrode impedance is then Zppy = Zserial/2. Similar procedure was repeated to obtain Zplatinum.

4.4. Cell-laden conductive scaffold platform

The Stem Cell Research Oversight committee at Stanford approved all stem cell procedures. Human induced pluripotent stem cells (iPSC) derived neural progenitor cells (hNPC) were cultured in a 6-well plate as described previously (11). Briefly described, iPSCs were exposed to specific culture conditions and growth factors (Noggin, bFGF, and retinoic acid) to steer their differentiation towards a neural fate. hNPCs of passages 2 to 3, were seeded at 100,000 cells/cm² on the PPY scaffold system with Matrigel as substrate matrix and cultured in NPC maintenance media. NPC maintenance media formulation was DMEM/F12 (50%), Neurobasal (50%), N2-MAX (1%), B27 (1%) non-essential amino acids (NEAA) (1%), GlutaMAX (1%), 2-mercaptoethanol (0.1 mM), penicillin/streptomycin (P/S, 1% v/v) supplemented with bFGF (20 ng/mL) and EGF (20 ng/mL). In-vitro assays were performed on day 3 with media change every day. In-vivo implantation was performed on day 2 after cell seeding.

4.5. In vitro cell viability assay

In vitro cell viability was assessed on day 3; one day after electrical stimulation (ES) was applied. Survival of cells after ES was determined by a live/dead staining kit (Life Technologies) as previously described (8). Briefly, the cells were incubated with 2 µL/mL of ethidium homodimer-1 and calcein AM for about 15 mins at 37 °C in the dark. The cells were then rinsed with 1X PBS and placed on a glass slide with coverslip. Cells on the scaffold were imaged using a fluorescent microscope (Keyence BZ-X700).

4.6. In vitro quantitative real-time polymerase chain reaction (qRT-PCR) analysis

RNA extraction and qRT-PCR analysis was performed as described previously (12). Total RNA was extracted from the cells seeded on the conductive scaffold 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). TaqMan-polymerase and primers (Qiagen, Germantown, MD) were used for qRT-PCR. Pre-developed TaqMan reagents used were human VEGFA (Hs00900055_m1), VEGFB (Hs00173634_m1), BDNF (Hs02718934_s1), NRN1 (Hs02786624_g1), and housekeeping genes - GAPDH (Hs02786624_g1). qRT-PCR was carried out on a QuantStudio 6 Flex Real-Time PCR System (Thermo Fisher, Waltham, MA). 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.

4.7. In vivo surgical procedures

All animal procedures were approved by Stanford University’s Administrative Panel on Laboratory Animal Care. Adult, male T-cell deficient rats (to reduce risk of immune rejection against transplanted human stem cells; NIH-RNU 230 ± 30 g) underwent distal middle cerebral artery (dMCA) occlusion model (permanent electrocoagulation) of stroke. Here, both common carotid arteries were also temporarily occluded lasting 30 min as previously described (12). Rats were anesthetized with 5% isoflurane and maintained on 2% isoflurane throughout the surgical procedure. Buprenorphine (0.05mg/kg) was administered subcutaneously for analgesia. Ampicillin was added to cage water 1 day prior to surgery (1mg/ml) and for 7 days after transplantation. Weight and pain symptoms of the animal were monitored continuously under supervision of veterinary staff for a week after surgery.

One week after stroke, animals were randomized by vibrissae-whisker paw score, and those with effective stroke (between 0–2 out of 10) were chosen for implantation. Implantation surgeries were performed by a blinded individual as described previously (12). Briefly, a craniectomy was drilled above the left cortical region between the neuroanatomical lambda and bregma markers, and the dural layer was excised. The conductive polymer scaffold (cultured with/without hNPC cells for 24 hours) was removed from the in vitro system and gently rinsed with 1xPBS. The scaffold was implanted over the exposed brain tissue primarily on the penumbral cortex medial to the lesion. A reference electrode attached to the wireless system was placed on the contralateral skull (Fig. 5b). The wireless chip was secured to the contralateral skull with dental cement and the external skin tissue was sutured and closed. Weight and pain symptoms of the animal were monitored continuously till recovery from surgery.

4.8. BrdU injections

BrdU intraperitoneal injections were performed 24 hours after the implantation of the conductive polymer system (post-stroke day 9). Briefly, BrdU was diluted in PBS to make a sterile solution of 10 mg/mL. A concentration of 100 mg/kg was intraperitoneally injected. Animals were sacrificed and fixed at 3 weeks post-stroke.

4.9. In vivo electrical stimulation

Rats in ES groups were electrically stimulated at −13dBm, 1 GHz for 1 hour. (Agilent N9310A) for the first 3 days after implantation (implantation on Day 8, electrical stimulation on Day 9, Day 10, and Day 11 post-stroke). The potential was applied across the implanted polymer and the reference electrode placed on the contralateral skull while the animals were freely moving within the cage wirelessly connected to stimulator system function generator (E3641A, Agilent). Rats were monitored continuously for normal behavior and pain symptoms throughout stimulation.

4.10. Stroke size assessment

Rats were perfused with 4% paraformaldehyde (Sigma Aldrich) at 3-week post stroke. The wireless scaffold system was removed from the euthanized animals prior to sectioning. About 40µm coronal slices were sectioned as described previously (12). Animal slices were used for stroke volume and immunostaining analysis. Stroke size was assessed using cresyl violet staining 3 weeks after stroke as described previously (24) Serial slices were taken 400 µm apart from the genu of the corpus callosum to the splenium. Brightfield images of the stained slices were imaged using Keyence All-in-One Fluorescence Microscope (BZ-X700, Keyence). Stroke size was calculated using ImageJ as per the following equation $\frac{Are{a}_{Contralateral}-Are{a}_{Ipsilateral}}{Are{a}_{Contralateral}}x100$. Assessments were performed by a blinded individual.

4.11. Immunohistochemistry

Serial slices were taken 400 µm apart from the genu of the corpus collosum to the splenium. Primary antibodies were incubated in a blocking buffer at 4°C overnight, followed by three 15-min PBS washes and detected by secondary antibodies (Alexa Flour 488, 555 or 647, Life Technologies) as described previously ⁶³. Primary antibodies include anti-BrdU (1:500, Abcam), anti-DCX (1:750, Abcam), HuNu (Anti Nuclei antibody, 1:250, Millipore). Samples were counterstained with DAPI (Hoechst 33342, 1:5000, Invitrogen) to visualize nuclei and mounted with fluoromount aqueous medium (Sigma-Aldrich) before imaging. Samples were imaged on a Keyence All-in-One Fluorescence Microscope (BZ-X700, Keyence) using 20X or 60X objectives and analyzed using ImageJ (NIH).

Cell imaging and counting were performed by a blinded individual. Peri-infarct area was defined as the area immediately surrounding the infarcted tissue and sampled in a 750µm x 650µm window. The SVZ region was defined in the same manner as previously (12) briefly tissue next to the ventricles with higher cellular density was assessed and 750µm×650µm frame was obtained using a Keyence All-in-One Fluorescence Microscope (BZ-X700, Keyence) using 20X objectives. For cell counting, the images were opened in ImageJ software with three color channels and grid view. Number of cells stained for BrdU, DCX and co-stained for both the stains were stereologically counted and averaged across slices for a given animal with n=4 per experimental group.

4.12. 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.