Tethering Peptides onto Biomimetic and Injectable Nanofiber Microspheres to Direct Cellular Response

Abstract

Biomimetic and injectable nanofiber microspheres (NMs) could be ideal candidate for minimally invasive tissue repair. Herein, we report a facile approach to fabricate peptide-tethered NMs by combining electrospinning, electrospraying, and surface conjugation techniques. The composition and size of NMs can be tuned by varying the processing parameters during the fabrication. Further, bone morphogenic protein-2 (BMP-2) and vascular endothelial growth factor (VEGF) mimicking peptides have been successfully tethered onto poly(ε-caprolactone) (PCL):gelatin:(gelatin-methacryloyl) (GelMA)(1:0.5:0.5) NMs through photocrosslinking of the methacrylic group in GelMA and octenyl alanine (OCTAL) in the modified peptides. The BMP-2-OCTAL peptide-tethered NMs significantly promote osteogenic differentiation of bone marrow-derived stem cells (BMSCs). Moreover, human umbilical vein endothelial cells (HUVECs) seeded on VEGF mimicking peptide QK-OCTAL-tethered NMs significantly up-regulated vascular-specific proteins, leading to microvascularization. The strategy developed in this work holds great potential in developing a biomimetic and injectable carrier to efficiently direct cellular response (Osteogenesis and Angiogenesis) for tissue repair.

Graphical Abstract

A facile approach for the fabrication of stable peptide decorated nanofibrous microsphere for the cellular responses like osteogenesis and angiogenesis. The dual peptide decorated nanofibrous microsphere would be an appealing choice for the tissue engineering applications.

Background

With applications as drug and cell delivery vehicles and cellular response regulators, microsphere development receives significant attention.^1–5 For decades, attempts to develop an ideal candidate with biomimetic nanofibrous morphology, injectability, and functionalization sites have fallen short.^6,7 However, several NMs with constrained compositions were generated using self-assembly.^8–12 Ma Laboratory pioneered nonfiber microsphere fabrication using self-assembly and explored their potential applications as cell and growth factor carriers.^10,11 Previously, Ma et al. used star-shaped poly(_L-lactic acid) (PLA) to fabricate hollow NMs via self-assembly to deliver chondrocytes for knee repair.¹⁰ In more recent studies, certain modifications of the material enabled controlled delivery of stem cells, growth factors, miRNA, and other biomolecules for tissue regeneration applications.^{7, 11, 12} Here, the composition and molecular weight (MW) of the polymers played a key role to achieve optimal fibrous morphology. Due to batch variability during polymer synthesis, ubiquitous MW and morphologies are difficult to achieve.⁹ Though necessary for surface functionalization, tethering functional groups like –NH₂, -COOH, and acrylic or methacrylic groups may inadvertently disrupt self-assembly and alter microparticle morphologies by changing polymer chain hydrophobicity.

Additionally, natural materials like collagen, gelatin, and chitosan can self-assemble into NMs.^{3, 6, 11} Zhang et al. reported that chitin chains rapidly self-assemble into nanofibers (NFs) in NaOH/urea aqueous solution, subsequently forming weaved microspheres through thermal induction.⁶ The same group then modified the chitin NMs with hydroxyapatite needles using a one-pot synthesized crystallization method to promote in vitro cell adhesion and in vivo bone healing.¹³ Recently, this group also fabricated cellulose NMs by dissolving cellulose in NaOH/thiourea aqueous solution and thermally-induced self-assembly.¹⁴ In a different study, Zhou et al. prepared extracellular matrix (ECM) mimicking chitosan NMs by physical gelation to promote chondrocyte seeding and proliferation, revealing chitosan NMs’ remarkable potential in cartilage tissue engineering.³ In contrast, Deng et al. constructed an ultralight, pure, natural aerogel by spraying segmented cellulose nanofibrils into liquid nitrogen and freeze-drying.¹⁵ Inspired by this work, our group combined electrospinning and electrospraying to develop a novel NM fabrication method and investigated the potential application for the NMs in cell delivery.¹⁶ However, incubation in culture media over a long period of time led to mechanical instability. Thermal crosslinking failed to achieve mechanical stability of NMs. Moreover, our NMs failed to incorporate various functional molecules, peptides or growth factors to simulate the biochemical cues in tissue repair and regeneration.

To facilitate tissue growth, cells are used with scaffolds or delivery vehicles. Due to their rapid proliferation, pluripotency, bioavailability, and immunomodulation, stem cells are an appealing choice to regenerate tissues.^17–19 Specifically, bone marrow-derived stem cells (BMSCs) attract warranted attention in regards to bone regeneration due to their active involvement in various bone-forming biochemical cues and role as osteoblastic progenitor cell sources.^{20, 21} To further enhance bone regeneration, locally sustained delivery of signaling molecules from carriers/scaffolds is a viable approach.^{22, 23} Osteoinductive biochemical factors, such as BMP-2 derived peptide KIPKASSVPTELSAISTLYL, play a pivotal role in bone cell signaling, migration, proliferation, differentiation, and maturation.²⁴ Controlled local delivery of osteoinductive biochemicals fails to achieve optimal efficacy in clinics due to lack of suitable implementation methods. Various immobilization strategies, including covalent bonding, physical adsorption, and entrapment, are used in various scaffolds to prolong component release.^{25, 26} Most immobilization strategies rely on adsorption is given the lack of bioconjugation sites in scaffolds. Further, weak interactions between scaffolds and signaling molecules often lead to a high initial burst rather than a long-term, sustained release. In comparison, covalent coupling strategies like carbodiimide²⁷ and photo-driven thiol-ene²⁸ acrylic²⁹ conjugations enable sustained release of peptides over prolonged periods.

Despite advancements, tissue engineering scaffolds still suffer from low clinical translatability due to low success in inducing angiogenesis, particularly post-implantation.^30,31 Because vascular endothelial growth factor (VEGF) and VEGF-mimicking materials play an important role, along with extracellular matrix (ECM), to facilitate microvascular network formation on scaffolds, their incorporation and delivery may prove key to increasing the practical application of tissue scaffolds in regenerative medicine.³¹ Interestingly, D’Andrea et al. designed a VEGF mimicking QK peptide that replicates the function of VEGF, inducing angiogenesis.³² Hence, QK immobilization to scaffolds is a promising strategy to enhance microvasculature formation and sustain local cellular functions necessary for repairing specific tissue defect sites.^33–35

Herein, we report a facile approach to fabricate peptide-tethered NMs by combining electrospinning, electrospraying, and surface conjugation techniques. NMs were crosslinked by vapor induction instead of thermal treatment, ensuring sufficient mechanical stability. Incorporating gelatin methacryloyl (GelMA) into fibers provided binding sites for covalent conjugation of biomolecules via photocrosslinking. The BMP-2-OCTAL functionalized NMs were used to examine the osteogenic differentiation of NM-seeded bone marrow-derived stem cells (BMSCs). Furthermore, QK-OCTAL conjugated NMs were used to investigate the microvascular formation of seeded HUVECs.

Methods

Fabrication of GelMA Containing NMs

The PCL:gelatin: GelMA NMs were fabricated following the protocols above-mentioned, but using PCL:gelatin:GelMA SFs (SFs) and 5% crosslinker. To conjugate BMP-2 and VEGF-mimicking QK peptides to GelMA-containing NMs, BMP-2 and QK peptides were modified with octenyl alanine amino acids to form BMP-2-OCTAL, and QK-OCTAL containing an acrylate group as a pendant at N terminus end (GenScript, Piscataway, NJ). The conjugation was conducted as follows. The customized BMP-2-OCTAL and QK-OCTAL peptides (100 μg/ml) and photoinitiator Irgacure 2959 (2 μg/ml) were dissolved in PBS and added to a vial with 2 mg GelMA NMs, followed by photocrosslinking at 365 nm for 20 mins. Microspheres were washed using deionized water and freeze-dried for further use. To quantify the content of tethered peptides and release kinetics, FITC and TRIC were conjugated to peptides to form BMP-2-OCTAL-FITC and QK-OCTAL-TRITC.

Osteogenic Differentiation of BMSCs on NMs

Around 2 mg PCL:gelatin:GleMA NMs with and without conjugation of BMP-2-OCTAL peptides were sterilized by soaking in ethanol overnight and UV exposure for 24 h. Then, microspheres were washed thrice in PBS and used for subsequent cell culture experiments. The cell adhesion and proliferation of BMSCs³⁶ on the microspheres were determined by co-culturing the microspheres with BMSCs on agar-coated (0.1 wt%) 24 well plates. The microspheres were dispersed in complete DMEM and uniformly distributed in the wells of agar-coated 24 well plates. Cells were seeded at a density of ~50,000 cells/well. After 24 h, microspheres were transferred to new agar-coated wells for culture. Then, NMs were cultured in the osteogenic differentiation medium, OsteoMax, for 7 and 14 days.

At certain time intervals, the microspheres were harvested by centrifugation at 1000 rpm for 1 min. The cell-seeded microspheres (with and without BMP-2-OCTAL conjugation) were washed thrice in PBS and fixed in 4% paraformaldehyde (PFA) at room temperature for 30 mins. Fixed microspheres were washed thrice in PBS and immersed in permeabilization and blocking buffer (1% bovine serum albumin and 0.1% Triton in PBS) for 1 h. To stain the osteogenic markers, appropriate dilution of the methanolic solution of Alexa Fluor 647 phalloidin conjugated Runx2, OCN and OPN antibodies (Invitrogen, Carlsbad, CA) in the blocking buffer was used and incubated with the fixed microspheres at room temperature in the dark overnight. The microspheres were washed with PBS and counterstained with DAPI. Images were acquired using a Zeiss LSM 710 confocal microscope equipped with excitation and emission filters.

HUVECs Culture and Microvascularization on NMs

Human umbilical vein endothelial cells (HUVECs) (Lonza, Walkersville, MD) were grown in endothelial basal medium (EBM-2 SingleQuots, Lonza) supplemented with 2% fetal bovine serum, hydrocortisone, hFGF-B, VEGF, R3-IGF-1, ascorbic acid, hEGF, GA-1000, and heparin. The cells were cultured in an incubator at 37 °C with 5% CO₂. Prior to seeding, ~2 mg PCL:gelatin: GleMA NMs with and without QK-OCTAL conjugation were sterilized by soaking in ethanol overnight and UV exposure for 24 h. The sterilized microspheres were washed thrice in PBS and used for cell culture experiments. The adhesion of HUVECs on microspheres was determined by co-culturing HUVECs with NMs on agar-coated (0.1 wt%) 24 well plates. Microspheres were dispersed in complete EBM-2 with above-mentioned supplements and uniformly distributed in the wells of agar-coated 24 well plates. Around 50,000 cells were seeded to each well. After 24 h, the microspheres were transferred to different agar-coated wells for incubation. The cells on microsphere were collected after incubation for 3 and 7 days and fixed with 4% paraformaldehyde at room temperature for 30 min. The fixed microspheres were washed thrice in PBS and immersed in permeabilization and blocking buffer (2% Bovine serum albumin and 0.1% Triton in PBS) for 1h. The cells stained with primary anti-human CD31 antibody (R&D Systems, Minneapolis, MN) an angiogenic marker according to the previous protocol. After washing with PBS, microspheres incubated with AlexaFluor 647 goat anti-mouse secondary antibody (1:1000) (Invitrogen). The cell nuclei were stained with DAPI. The stained HUVECs were observed using a Zeiss LSM 710 confocal microscope equipped with appropriate excitation and emission filters.

Results

Figure 1 shows a illustrates the fabrication process of peptides-tethered NMs for regulating cellular response as follows: i) Cryocut or homogenize electrospun NF mats to generate NF segments; ii) Add crosslinker (e.g., gelatin) to the NF segment solution and homogenize the solution; iii) Electrospray the crosslinker-containing NF segment solution into liquid nitrogen; iv) Freeze-drying the obtained microspheres; v) Crosslink the microspheres after freeze-drying; and vi) Tether peptides to the microspheres using different surface conjugation techniques for regulating cell response such as osteogenesis and angiogenesis.

Figure 1.

Scheme illustrating the fabrication of peptides-tethered NMs and their applications. (i) Cryocut or homogenize electrospun nanofiber mats to generate nanofiber segments. (ii) Add crosslinker (e.g., gelatin) to the nanofiber segment solution and homogenize the solution. (iii) Electrospray the crosslinker-containing nanofiber segment solution into liquid nitrogen. (iv) Freeze-drying the obtained microspheres. (v) Crosslink the microspheres after freeze-drying. (vi) Tether peptides to the microspheres using different surface conjugation techniques for directing cellular response such as osteogenesis and angiogenesis.

Preparation and Characterization of Short Electrospun Nanofibers

Short electrospun NF segments were prepared by cryocutting and homogenizing with a probe sonicator. The length of segmented NFs was characterized by confocal laser scanning microscopy (CLSM). Aligned PCL:gelatin (1:1) and PCL:gelatin:GelMA (1:0.5:0.5) NF mats were cryocut to segments of 20 – 40 μm in length. Figure S1 shows confocal microscopy images of the homogenized PCL:gelatin and PCL:gelatin:GelMA SFs from 20 mg/ml dispersion, showing uniform dispersion in water. PCL possesses a low glass transition temperature of −50 °C, hence it cannot fragment by homogenization due to its viscoelastic properties. Therefore, PCL:gelatin and PCL:gelatin:GelMA fibers were segmented into SFs by cryocutting, and dispersed in water by homogenization.

Engineering PCL:gelatin NMs

To avoid thermal crosslinking, we added gelatin as a crosslinker during the homogenization for glutaraldehyde (GA) vapor-induced crosslinking. GA-induced crosslinking is widely used to increase the mechanical stability of NF mats.³⁷ Further, we employed all other fabrication techniques from our previous work.¹⁶ In the current study, parameters like applied voltage (4–6 kV), the flow rate of fiber dispersion (2.0 ml/h), and distance (10 cm) between the spinneret and freezing media were fixed for all fabrication processes. Figure S2 shows the morphology of PCL:gelatin (1:1) NMs with crosslinker ratios ranging from 20%, 10%, to 5% with respect to the fiber content in dispersion. SEM images show mostly uniform particle size in each composition. We can control the size of NMs by adjusting flow rate and/or applied voltage, which was reported in our previous work.¹⁶ Figure S2(a-d) shows the PCL:gelatin NMs with 20% crosslinker ratio before and after crosslinking. NM shrinkage was observed after crosslinking with GA. After crosslinking, the shape of the particles varied from spherical to irregular, due to the aldehyde amine covalent bonding. Figure S2(e-h) shows SEM images of PCL:gelatin NMs with 10% crosslinker before and after crosslinking. A similar shrinking effect was observed for NMs with 10% crosslinker ratio. Figure S2(i-l) shows SEM images of PCL:gelatin NMs with 5% crosslinker ratio before and after crosslinking. Interestingly, no particles shrunk after crosslinking with GA. To further verify the crosslinking influence of gelatin with GA, we prepared PCL:gelatin (1:1) NF particles with 15 mg/ml fiber dispersion in water with different crosslinker ratios ranging from 20%, 10% to 5%. A similar shrinking scenario for the NMs was observed after crosslinking with GA vapor which was fabricated with 20% (Figure S3(c-d)) and 10% (Figure S3(g-h)) crosslinker ratios. Similarly, there is no significant shrinkage observed for NMs generated with 5% crosslinker ratio (Figure S3(i-l)). These results demonstrate that NMs fabricated with 5% crosslinker ratio can maintain shape and mechanical stability, as compared to thermally crosslinked NMs.¹⁶ In addition, well-dispersed crosslinkers within NMs provide additional support to make a dynamic bond between the fiber segments which hold each other, thus maintaining morphology during fabrication. Hence, MNs composed of different inorganic and organic polymer SFs could be readily engineered using this approach.

Fabrication of GelMA Containing Nanofiber Microspheres

Given its biocompatibility and biodegradability, PCL is a popular material for fabricating tissue engineering scaffolds.^{38, 39} However, PCL’s lack of functionality sites poses a challenge for conjugating various signaling molecules for enhancing functionality in biomedical applications. PCL:gelatin NF microparticles have similar issues. Moreover, the majority of surface conjugations are involved with solution-based chemistries. Resultantly, we prepared PCL:gelatin:GelMA (1:0.5:0.5) NMs for photo-driven conjugation of peptides and biomolecules based on methacrylate functionality. Figure 2 shows the SEM image of the PCL:gelatin:GelMA NMs. Particles exhibit relative uniformity, with sizes 300–400 μm. Additionally, no shrinkage was observed after GA crosslinking for 12 h (Figure 2(c–d)). These NMs were used for further studies, including peptide conjugation and cell seeding.

Figure 2.

SEM images showing NMs composed of PCL:gelatin:GelMA (1:0.5:0.5) nanofiber segments (a, b) before and (c, d) after crosslinking. The nanofiber segment solution contained 5% gelatin as crosslinker during the electrospraying process.

Photo-driven Conjugation of Peptides to Nanofiber Microspheres and Release Kinetics

Given its ability to immobilize peptides onto polymer substrates, photo-crosslinking as a conjugation strategy is broadly utilized.^{40, 41} Here, we customized BMP-2 peptides and VEGF-mimicking QK peptides with OCTAL to introduce allyl functionality in the peptide chain for the conjugation with NMs. The fluorescein isothiocyanate (FITC) and tetramethylrhodamine (TRITC)-coupled peptides (like BMP-2-OCTAL-FITC and QK-OCTAL-TRITC) were used to quantify NM-conjugated peptides. Figure 3(a) illustrates BMP-2-OCTAL and QK-OCTAL conjugation to PCL:gelatin:GelMA NMs under UV light (365 nm) in the presence of photoinitiator Irgacure 2959 for 20 mins. Figure 3(b) shows confocal microscopy images of PCL:gelatin:GelMA NMs before and after BMP-2-OCTAL-FITC conjugation. Figure 3(b)(i-iii) shows the PCL:gelatin:GelMA microparticles without conjugation of BMP-2-OCTAL-FITC. The green color with low intensity was due to the autofluorescence during the crosslinking of GA vapor with gelatin. In contrast, BMP-2-OCTAL-FITC conjugated NF microparticles showed very high fluorescence intensity (Figure 3(b)(iv-vi)). Further demonstrating its versatility, QK-OCTAL-TRITC was conjugated to PCL:gelatin:GelMA NMs in the presence of a photoinitiator under UV light irradiation at 365 nm for 20 mins. Figure 3(c) shows confocal microscopy images of PCL:gelatin:GelMA NMs before and after QK-OCTAL-TRITC conjugation. The NMs after conjugation fluorescent red with a significantly higher intensity as compared to the non-conjugated NMs (Figure 3(c)(i-vi)). The conjugation occurred GelMA moiety (short fiber) in both surface and core of microspheres, and the morphology remained the same after the conjugation of peptides.

Figure 3.

(a) Scheme illustrating the photo-driven conjugation between methacrylic group containing NMs and OCTAL modified peptides. (b) Confocal microscopy images of PCL:gelatin:GelMA (1:0.5:0.5) NMs (i)-(iii) before conjugation of BMP-2-OCTAL-FITC (showing some autofluorescence) and (v)-(vi) after conjugation of BMP-2-OCTAL-FITC. (c) Confocal microscopy images of PCL:gelatin:GelMA (1:0.5:0.5) NMs (i)-(iii) before conjugation of QK-OCTAL-TRITC and (v)-(vi) after conjugation of QK-OCTAL-TRITC. (i, iv): fluorescent images. (ii, v) optical images. (iii, vi) merged images.

To quantify BMP-2-OCTAL conjugation, BMP-2-OCTAL-FITC was used to conjugate with NMs in presence of the photo initiator I2959 by UV light at 365 nm for 20 mins. Roughly 48±3% of BMP-2-OCTAL was successfully conjugated with NMs within 20 mins, while only 32±5% of BMP-2-OCTAL was adsorbed after 24 h without using photoinitiation. The release kinetics of BMP-2 peptides after conjugation to NMs was shown in Figure 4. The initial burst release was observed from the BMP-2-OCTAL-FITC incorporated NMs, which could be due to the absorbed BMP-2-OCTAL-FITC. The released peptides were more than 5 μg/ml at each time point in the first three weeks. The released peptides still reached more than 1 μg/ml even at the final day of release test. The results suggest that conjugated NMs can efficiently release BMP-2 peptides in a controlled and sustained manner. Similarly, QK-OCTAL-TRITC peptides were conjugated to the NMs. Roughly 46±4% of QK-OCTAL-TRITC peptides was conjugated to 2 mg of PCL:gelatin:GelMA NMs. Given conjugation strategy and base material similarity, release kinetics of QK-OCTAL-TRITC peptides matched that of BMP-2-OCTAL-FITC.

Figure 4.

In vitro release curves of BMP-2-OCTAL-FITC from BMP-2-OCTAL-FITC conjugated PCL:gelatin:GelMA (1:0.5:0.5) NMs over a period of 30 days. Error bars represent means ± SD (n = 3).

Biomineralization of Nanofiber Microspheres

Fabricated PCL:gelatin:GelMA NMs with and without BMP-2-OCTAL conjugation were immersed in a simulated body fluid (SBF) to evaluate their bioactivity in vitro. Calcium apatite formation on the NF particles indicates bioactivity.⁴² After 12 h of incubation in SBF, deposited minerals were observed on both BMP-2-OCTAL conjugated and non-conjugated NMs. SEM images show the morphologies of deposited apatite-like minerals (Figure S4). Figure S4(a-c) and Figure S4(d-f) show the apatite-like mineral deposition on PCL:gelatin:GelMA NMs without and with BMP-2-OCTAL peptide conjugation, respectively. The elements of formed apatite-like minerals on the NMs were identified by EDX analysis, indicating the deposition of a significant amount of calcium and phosphorus.

Osteogenic Differentiation of BMSCs Seeded on NMs

In a previous sudy, we observed BMSCs proliferation on NMs.¹⁵ Here, an increased seeding ratio of BMSCs (~5 × 10⁴ cells) was used for 2 mg PCL:gelatin:GelMA NMs without and with the conjugation of BMP-2-OCTAL. The ECM-mimicking nature of NMs provides an exceptional microenvironment for both cell adhesion and proliferation. BMSCs proliferation on the NMs was reported in our previous work, ascertaining our observations in this study.¹⁶ Here, we focused on the investigation of osteogenic differentiation of BMSCs on the NMs with and without BMP-2-OCTAL conjugation.

BMSCs seeded on the NMs with and without BMP2-OCTAL conjugation were cultured in differentiation medium for 7 and 14 days. At each time point, cells were stained with specific osteogenic markers Runx2, OCN, and OPN with corresponding secondary antibodies. Figure S5 shows confocal microscopy images of Runx2 expression (pink color) of BMSCs seeded on the NMs without and with BMP-2-OCTAL conjugation. Evidently, Runx2, an early osteogenic marker, was clearly visualized at day 7 and it started fading at day 14 in both groups. However, higher Runx2 expression was observed in BMSCs seeded on BMP-2-OCTAL-tethered NMs compared to cells seeded on non-conjugated ones. In addition, immunostaining of late osteogenic markers like OPN and OCN was also performed as shown in Figure 5, indicating BMSCs seeded on BMP-2-OCTAL-conjugated NMs had higher OPN expression (pink color) at day 14 than the cells seeded on non-conjugated particles. Figure S6 shows the staining of OCN expression in BMSCs which were seeded on NMs with and without conjugation of BMP-2-OCTAL. The BMSCs seeded on BMP-2-OCTAL-conjugated NMs had higher OCN expression (pink color) at day 14 than cells seeded on the microparticles without tethering peptides. Since OPN and OCN are late osteogenic markers, they were not significantly presented on day 7.

Figure 5.

Confocal microscopy images showing OPN expression of BMSCs seeded on PCL:gelatin:GelMA (1:0.5:0.5) NMs without (a) and with (b) conjugation of BMP-2 peptides after culturing in the osteogenic differentiation medium for 7 and 14 days. The cells stained with OPN antibody in pink and counterstained with DAPI in blue. The right column shows merged images of fluorescent images of OPN antibody staining and DAPI staining and the bright field images.

To further examine the bioactivity of BMP-2-OCTAL-tethered peptides on BMSCs osteogenic differentiation, gene expression was analyzed by reverse transcription-polymerase chain reaction (RT-PCR) after 7 and 14 days. BMSCs seeded on the particles without BMP-2-OCTAL conjugation were used as a control. Osteogenic-specific genes (OCN, OPN, and Runx2) and related matrix gene ALP were assessed. Figure 6 shows the osteogenic gene expression. The expressions of ALP, OPN, and OCN were higher for BMSCs seeded on BMP-2-OCTAL conjugated NMs than the control at both day 7 and 14. Further, BMSCs seeded on BMP-2-OCTAL-conjugated NMs had higher Runx2 expression than the control at day 14, though no significant difference was observed at day 7. Gene expression results were reaffirmed with immunohistochemistry analysis. The overall results demonstrated that BMP-2-OCTAL-conjugated NMs promote osteogenesis in vitro compared to the particles without peptide conjugation, indicating that the BMP-2 peptides following OCTAL conjugation maintain biological functions.

Figure 6.

mRNA expression of osteogenic-specific genes and related matrix genes including (a) ALP, (b) OPN, (c) OCN, and (d) Runx2 of BMSCs cultured on PCL:gelatin:GelMA (1:0.5:0.5) NMs with and without (control) conjugation of BMP-2 peptides for 7 and 14 days. *p < 0.05

Microvascular Networks Formation on NMs

VEGF is a proven potent angiogenic factor, playing a critical role in the formation of vasculature structures.^{31, 32} QK peptides consisting of 15 amino acids have been designed to recognize VEGF receptors, regulating both endothelial cell proliferation and propensity toward angiogenesis.³² In this work, OCTAL tethered QK (QK-OCTAL) peptides were conjugated with NMs to induce microvascularisation of microsphere-seeded HUVEC cells. Around 5 × 10⁴ cells were seeded to 2 mg of PCL:gelatin:GelMA NMs with and without QK-OCTAL conjugation. Figure 7 shows CD31, an endothelial cell marker, (in pink) and DAPI (in blue) staining of HUVECs on NMs with and without QK-OCTAL conjugation after incubation for 3 and 7 days. Though no significant tube formation was observed on either the control or QK-OCTAL-conjugated NMs at day 3, endothelial tubule roots were observed by day 3 on QK-OCTAL-conjugated NMs, as indicated by the red arrows (Figure 7b). Interestingly, dense microvascular networks were formed on QK-OCTAL-conjugated NMs after incubation for 7 days. Additionally, QK-OCTAL-conjugated NMs showed denser, mature, and interconnected microvascular networks (longer tubes) compared to NMs without tethered QK peptides. The number of nodes and vascular tubes was quantified as illustrated in Figure 8. Around 50 nodes per view existed on QK peptide-conjugated NMs, while fewer than 10 per view existed on control NMs following incubation for 7 days (Figure 8a). Similarly, more than 100 microvascular tubes per view on the QK peptide-conjugated NMs; however, there were only 30 microvascular tubes per view on the control NMs (Figure 8b).

Figure 7.

Confocal microscopy images showing tubular network formation from HUVECs seeded on PCL:gelatin:GelMA (1:0.5:0.5) NMs without (a) and with (b) conjugation of QK peptides in the simulated medium for 3 and 7 days. The cells stained with CD31 in pink and counterstained with DAPI in blue. The right column shows the merged images of fluorescent images of CD31 staining and DAPI staining and the bright field images.

Figure 8.

Quantification of nodes and vascular tubes on the NMs with and without (control) conjugation of QK peptides. (a) Number of nodes/microsphere after incubation of HUVECs for 3 and 7 days. (b) Number of vascular tubes/microsphere after incubation of HUVECs for 3 and 7 days. *p < 0.05

Discussion

Direct injection of cells to diseased sites demonstrated some success in modern medicine. However, many issues including the low survival, and no cellular integration remain unmet.⁴³ To address these issues, cell-laden 3D injectable hydrogels and NMs have been developed.⁴⁴ Likewise, injectable biomaterials with suitable functionalization sites show great promise as delivery vehicles for fast regeneration of diseased or injured tissues.⁸ In this work, we developed peptide-tethered biomimetic and injectable NMs to direct cellular response. Micron-scale NF spheres synthesized in this study hold several unique features: i) ECM biomimetic (nanofibrous morphology), ii) ability to fill irregularly-shaped defects (micron size), iii) forming porous structures after injecting/packing to the defect site with good interconnectivity (geometric shape), vi) exerting biochemical cues to direct cellular response (tethered peptides).⁴⁵

Studies devoted to fabricating injectable NMs utilizing a variety of polymers, including star-shaped poly(_L-lactic acid) (ss-PLLA), polyhydroxybutyrate (PHB), chitin, and cellulose via phase separation and self-assembly, have been previously investigated.^{8, 46} However, such studies fail to facilitate precise control over particle size, are limited to certain compositions, and prove difficulty in large-scale production. Recently, we developed a novel NMs synthesis strategy by electrospraying aqueous dispersions of short, electrospun NF segments into a cryo-coolant, followed by freeze-drying and thermal crosslinking, thus producing NMs with a broad range of polymer compositions, different morphologies, and different sizes with narrow size distribution.¹⁶ Our fabricated NMs improved biological function of BMSCs and neural differentiation of embryonic stem cells, as compared to solid NMs.

Certain limitations in using electrosprayed microparticles, such as insufficient mechanical stability for the long-term culture, challenge the use of such microparticles in further studies. In our previous study, thermal crosslinking at ~50 °C successfully strengthened the particles, however, the melting points of both PCL and gelatin are close to that value.¹⁶ Therefore, crosslinking at 50 °C may alter the nanofibrous morphology of NMs. Among crosslinking strategies, GA vapor-induced crosslinking is widely used for biomaterials like collagen, gelatin, and hyaluronic acid, due to the ease of control and prevention of fiber matrix collapse during crosslinking in an aqueous environment.⁴⁷ Moreover, GA vapor-induced crosslinking significantly increases material mechanical properties.⁴⁸ In this study, GA vapor-induced crosslinking between fiber surface and crosslinker gelatin maintained nanofibrous morphology, resulting in sufficient mechanical stability for the long-term cultures. Compared to our previous fabrication technique, varying amounts of crosslinker gelatin were incorporated into SF solutions during the homogenization including 20%, 10%, and 5% with respect to the SF concentration. NMs with 20% and 10% crosslinkers shrunk from spheres to irregular shapes following GA-vapor crosslinking, a phenomenon that can be attributed to the high content of crosslinkers in NMs. In contrast, NMs with 5% crosslinker retained spherical shapes after crosslinking, and the stability of particles was enhanced, as compared to thermal crosslinking ones. Moreover, other green cross-linkers like genipin and tannic acid will be used to replace GA in our future studies, if there is toxicity in animal studies.^49,50

The cellular response is a dynamic process largely regulated by various growth factors and proteins in the ECM. Growth factor-mimicking peptides conjugated to injectable NMs may direct local biochemical cues to manipulate seeded cell activity, which can enhance tissue regeneration and repair.⁸ QK peptides modified with single acrylate (OCTAL) group and multiple acrylate groups in previous studies without losing any significant bioactivity.^34,35 Here, a methacrylate group bearing PCL:gelatin:GelMA NMs with methacrylate functional sites was used to tether the ally bearing peptides like BMP-2-OCTAL and QK-OCTAL using photo-driven conjugation. Similar photochemistry is widely used to prepare hydrogels in gelation for tissue engineering.^40,41 To quantify the conjugation and release of BMP-2-OCTAL and QK-OCTAL peptides, the fluorophore-containing peptides including BMP-2-OCTAL-FITC and QK-OCTAL-TRITC were used to conjugate with NMs. The conjugation efficiency of BMP-2 and QK peptides to NMs under UV light in the presence of the photo initiator could reach 48±3% and 46±4% within 20 mins. In contrast, only 32±4% peptides were physically adsorbed to NMs even in 24 h. The conjugation density of the peptide can be tuned by changing the GelMA concentrations during the fabrication.

Here, NMs conjugated with BMP-2-OCTAL peptide were used to quantify the release kinetics. An initial burst release was noted, which is attributed to the non-conjugated, physically trapped peptides. Since the peptide conjugation occurred on the GelMA methacrylate moiety, the release profiles mainly depended upon the degradation of GelMA. Especially the cleavage of a covalent bond between the gelatin and BMP-2 conjugated methacrylate moiety contributed to the sustained release of peptides at later stages. Because of the same conjugation method and the similar molecular weights of BMP-2 and QK peptides, the release kinetics of QK peptides were assumed similar to BMP-2 peptides and not performed from QK-OCTAL conjugated NMs. In the future, we will attempt to adjust crosslinking times to tune the release profiles.

BMP-2 mimicking peptides enhance osteogenic differentiation of BMSCs, making them favorable to deliver for bone regeneration.²⁴ Therefore, BMP-2 peptide-conjugated NMs promote osteogenic differentiation of BMSCs seeded on the surface more than NMs without peptide conjugation. BMP-2 mimicking peptides exhibit good stability after conjugation to NMs. Tethered BMP-2 peptides induced higher BMSC expression of Runx2, OCN and OPN from BMSCs, the principal osteogenic markers during bone formation. Resultant osteogeneic gene expression of ALP, OPN, Runx2, and OCN are consistent with immunostaining results, suggesting that BMP-2 peptides conjugated NMs promote osteogenic differentiation of BMSCs at both genetic and protein levels, which further confirmed the bioactivity of immobilized peptides. Moreover, the special active sites of various amino acids including Asp (aspartic acid) and phosphorylated Ser (serine) from conjugated peptides could enhance the mineralization of the natural bone matrix.⁴⁸ Our biomineralization results show more calcium apatite deposition on BMP-2 peptides-tethered NMs as compared with the control.

Tissue regeneration is a complex process with contingent on forming vascularised networks that supply oxygen and nutrients to engineered tissues.^30,31 Particularly, microvascular formation is vital in bone regeneration, as new blood vessels not only provide oxygen and nutrients to highly metabolically-active regenerating calluses, but also present a route for inflammatory cells and bone precursor cells to reach the injury site.^51–53 VEGF secreted by inflammatory cells and stromal cells has been demonstrated to stimulate the formation of new blood vessels in the tissue defect area, thus offering a method of enhancing vascular formation.⁵² With this in mind, we conjugated VEGF-mimicking peptide QK-OCTAL to NMs. As demonstrated, QK-OCTAL peptide-conjugated NMs initiated tubular formation at day 3, forming mature microvascular networks on microsphere surface at day 7. The statistical analysis revealed that higher numbers of junction nodes and vascular tubes were formed on QK-OCTAL conjugated NMs as compared with the control.

Conclusion

In summary, we have demonstrated a facile and generic approach for the fabrication of biomimetic and injectable peptide-tethered NMs by combining electrospinning, electrospraying, and surface conjugation techniques. BMP-2 peptide-conjugated NMs promoted osteogenic differentiation of BMSCs at both the genetic and protein levels. Further, a significantly higher number of junction nodes and vascular tubes were formed from HUVECs on VEGF-mimicking QK peptide-conjugated NMs compared to non-conjugated particles. The manufacturing approach presented in this work is a platform technology with the flexibility to tether a wide variety of peptides/biomolecules to NMs for directing cellular response. The developed NMs can be used as advanced and versatile injectable carriers for cell delivery to diseased tissues or tissue defects through a minimally invasive procedure. Financial Support Information: This work was supported by grants from the National Institute of General Medical Science (NIGMS) at the NIH (2P20 GM103480), National Institute of Dental and Craniofacial Research (NIDCR) at the NIH (1R21DE027516), NE LB606, and startup funds from the University of Nebraska Medical Center.