Abstract
Bioadhesive materials and patches are promising alternatives to surgical sutures and staples. However, many existing bioadhesives do not meet the functional requirements of current surgical procedures and interventions. Here we present a translational patch material that exhibits: (1) instant adhesion to wet tissues (2.5-fold stronger than Tisseel, an FDA-approved fibrin glue), (2) ultra-stretchability (stretching to >300% its original length without losing elasticity), (3) compatibility with rapid photo-projection (<2 min fabrication time/patch), and (4) ability to deliver therapeutics. Using our established procedures for the in silico design and optimization of anisotropic-auxetic patches, we create next generation patches for instant attachment to wet and dry tissues while conforming to a broad range of organ mechanics ex vivo and in vivo. Patches coated with exosomes demonstrate robust wound healing capability in vivo without inducing a foreign body response and without the need for patch removal that can cause pain and bleeding. We further demonstrate a new single material-based, void-filling auxetic patch designed for the treatment of lung puncture wounds.
Teaser We demonstrate a sticky and highly elastic patch with conforming designs for dynamic organ repair.
Introduction
Bioadhesive materials and patches have attracted considerable academic and commercial interest as replacement materials for sutures, staples, and wound dressings (1–3). While sutures and staples remain the current standard approach for closing wounds and cuts, they do not always prevent leakage and can contribute to surgical wound complications, toxicity (4), inflammatory responses (5), undesirable matrix remodeling and scar tissue formation (5), and post-operative pain (6). While staples are a valid alternative to sutures and can be used in mass casualty situations due to their quick application, they are not applicable in every situation. For instance, they cannot usually be used to close wounds on the hands and feet because they can cause discomfort (7).
These limitations of current approaches have prompted interest in using bioadhesive materials and patches for wound closure, wound healing, air leak management, and tissue sealing. However, many existing bioadhesives do not meet the functional requirements of modern surgical procedures and interventions. For example, liquids and glues are easily displaced and diluted when applied to wet tissues and/or internal organs. Other bioadhesives become ineffective and completely lose their tissue adhesiveness in wet environments or when they encounter bodily fluids. Many other bioadhesive materials require at least several minutes of tight tissue-patch interfacing for effective tissue adhesion (8, 9), making them unsuitable for attachment to dynamic organs in the clinical setting.
Additional barriers to the clinical translation of bioadhesive materials include insufficient flexibility and an inability to conform to the mechanics of internal organs and the skin. Dynamic organs including the lungs and heart present specific challenges towards the effective use of tissue adhesives and patch-based therapeutics (1). Their repetitive physiological motion and volumetric expansion and contraction create a fatigue loading environment that predisposes patches or bioadhesives to premature breakage or detachment (1, 10). Patches are typically held in place using sutures, staples, or surgical glues that can similarly detach or strain the organ over time, require specialist skills to apply, and present post-operative complications (inflammation, scarring, etc. (5)). For external skin applications, it is crucial that the bioadhesives and patches provide sufficient flexibility and elasticity to withstand tension caused by twisting and bending and remain attached. Furthermore, to avoid discomfort and pain from repeated dressing changes, patches for skin wound healing should be biodegradable and disintegrate without causing a foreign body response.
In many applications, applying a bioadhesive patch alone does not promote tissue healing, so delivering therapeutics and drugs via bioadhesive materials and patches would be desirable to accelerate the healing process. Integration of therapeutics and drugs into liquids, glues, and hydrogels can, however, be challenging due to their inability to retain the drugs. Thus, an ideal bioadhesive or patch should be: (1) biocompatible and biodegradable, (2) strongly and rapidly adherent, (3) elastic and mechanically conform to the underlying tissue, and (4) able to deliver therapeutics to accelerate healing.
In our previous work, we established a new design framework for developing hydrogel patches with anisotropic and auxetic properties (11). However, this initial set of materials had limited stretchability and were not intrinsically tissue adhesive. Attaching the patches to tissues required the application of fibrin glue, which is expensive so precludes their use in low-resource settings.
Therefore, the objective of this work was to develop the next generation of patch materials with the following characteristics: (i) instantaneous and strong tissue adhesion; (ii) ultra-elasticity, (iii) adaptable and scalable printing into precise anisotropic-auxetic architectures for a wide range of organs and tissues, (iv) biocompatibility and biodegradability, and (v) ability to deliver therapeutics. We applied our previously established framework for creating anisotropic-auxetic architectures (11) to develop an ultra-elastic and instantly adhesive patch platform adaptable to a broad range of tissue-specific mechanical properties, including elasticity, stiffness, and rigidity. Here, we demonstrate how our new patch system can be readily fine-tuned for instantaneous attachment to different organs and conform to organ mechanics ex vivo and in vivo. Our patches show substantially stronger attachment to wet tissues compared with commercially available fibrin-based Tisseel® (Baxter). Lastly, we evaluate the therapeutic potential of the patches in two exemplar pathologies, wound healing and pulmonary air leakage. We demonstrate that the adhesiveness of the patches allows coating with exosomes for enhanced wound healing without the need for patch removal and without inducing a foreign body response. With respect to pulmonary air leakage, we demonstrate how a single patch material can be fine-tuned to prevent air leakage from a punctured lung while conforming to the lung’s auxetic mechanics.
Results
AuxES patches - material composition and patch fabrication
Our platform of auxetic, elastic, and sticky (AuxES) patches, is based on a novel composition of materials that when combined lead to its unique elasticity, adhesiveness, and high-resolution lattice structure. Notably, the materials on their own, including gelatin methacryloyl (GelMA), acrylic acid (ACA), calcium chloride (CaCl2), and curcumin nanoparticles (CNPs) do not deliver the desired characteristics, but only when combined are able to conform to dynamic organ mechanics and achieve therapeutic functionality. The patch fabrication procedure shown in Figure 1A is simple, cost effective, rapid, and highly reproducible. The material can easily be formulated using off-the-shelf reagents (all reagents can be commercially procured) and is selectively photocrosslinkable, rendering it compatible with digital light projection (DLP), which has specific advantages over other conventional or advanced manufacturing techniques (12). In DLP, a digital micromirror device is used to project the image of a construct one layer at a time, which allows facile fabrication of highly complex architectures, which is otherwise challenging using conventional techniques such as casting or electrospinning (12).
A. The AuxES patches are fabricated via a facile single-layer projection-printing of a photoink (GelMA+ACA) contained in a PDMS-coated substrate. After washing and neutralization, calcium chloride is used to enhance patch integrity and adhesion. B. Optimization of CNP content within AuxES patches for high-resolution printing. Since CNPs act as a radical quenching agent, increasing their concentration improves print resolution. However, it also reduces the UV penetration depth within patches (for the same UV exposure duration). A 2 mg/ml CNP concentration was chosen since it results in high print resolution (∼375 μm minimum feature size) and consistent thickness of ∼1.2 mm. C. The patch materials demonstrate excellent biocompatibility with or without CNP as tested with viability of 3T3 cells after 48 h in culture.
We chose gelatin as the polymer backbone for the matrix as it is widely studied and has good biocompatibility, possesses integrin binding sites (RGD) for cell adhesion, and can undergo enzyme-mediated biodegradation (13–15). GelMA is prepared through the methacrylation of gelatin and can undergo rapid chain growth polymerization in the presence of a photoinitiator, providing excellent compatibility with high-resolution photocrosslinking (16, 17). GelMA has become the gold-standard for bioprinting in recent years due to its easy and replicable production, however on its own does not provide tissue adhesion (17, 18). We used LAP (lithium-phenyl-2,4,6-trimethylbenzoylphosphinate) photoinitiator (0.03% w/v) due to its compatibility with visible UV light (405 nm) crosslinking and biocompatibility (14, 19). The addition of ACA did not affect the photocrosslinking of the matrix formulation, and the use of a collimated light beam in DLP allowed patch fabrication as a single layer in under 2 minutes (18), substantially faster than extrusion printing (12).
CNPs act as a UV-absorptive and radical quenching agent critical for high-resolution DLP fabrication (18). We optimized the matrix CNP concentration by determining the print resolution and thickness of template patches with a broad range of feature sizes (Figure 1B) made using different CNP concentrations. Increasing the CNP concentration significantly reduced the minimum achievable feature size (p < 0.05) but also reduced patch thickness by affecting the UV penetration depth. We finally selected 2 mg/ml CNP, since it achieved a minimum feature size of 375 μm and a consistent patch thickness of ∼1.2 mm, which produced stable patches (Figure 1B).
For any new implantable medical device, the demonstration of lack of cytotoxicity is an essential first step. We therefore evaluated the viability of 3T3 fibroblasts cultured for over 48 h in the presence of the patches. In addition to CNP-laden patches, we also tested CNP-free patches made using FD&C yellow food dye, which has previously been shown to demonstrate the radical quenching needed for high-resolution DLP (18). The presence or absence of CNP had no effect on cell viability (∼100% after 48 h, Figure 1C).
Patches demonstrate instantaneous and strong bioadhesion to wet tissues
A material with strong and instantaneous bioadhesion can substantially reduce the time needed for patch application, thereby improving acceptance from surgeons and minimizing intraoperative harm (1). Attachment tests on wet organs (Figure 2A) using our AuxES patch formulation demonstrated strong and instantaneous (<1 s) bioadhesion to wet organs and could support the entire weight of a mouse liver upon immediate contact (see Supplementary Video S1). We further measured the forces required to detach the patches attached to mouse livers (Figure 2B), and the adhesion force for AuxES patches was almost twice that (p < 0.05) of commercially available fibrin glue (TISSEEL, Baxter Healthcare) and ∼10-times that of GelMA patches. When stretched, AuxES patches broke but did not detach from the organ (Supplementary Video S2), further demonstrating the strength of bioadhesion. Importantly, the AuxES patches not only strongly adhered to tissues but also imparted the ultra-elasticity needed to conform to various dynamic organs (Supplementary Video S3). The strong bioadhesion of the patches can be attributed to a combination of hydrogen and ionic bonding between the patch and the underlying tissue matrix (Figure 2C). Addition of acrylic acid to GelMA increases the number of carboxylic acid groups in the photocrosslinked matrix (20), which enhances patch bioadhesion to wet tissue through absorption of interfacial water and hydrogen bond formation, while calcium chloride results in ionic crosslinking of patches with the carboxylic acid groups in tissue matrix to further improve bioadhesion. The importance of CaCl2 treatment post patch printing is also highlighted in Supplementary Video S2, where patches without Ca2+ crosslinking are more easily detached from the organ, while patches with Ca2+ crosslinking do not detach from the organ. The prevalence of ionic and hydrogen bonding, in addition of the covalent binding of methacrylate moieties (Figure 2C), also improves the stretchability of the matrix (Figure 2A, Supplementary Video S3), which is discussed in further detail in the subsequent section.
A. Patches attached to a healthy mouse liver (removed from mouse cadavers) demonstrate instantaneous adhesion in <1 s sufficient to support the weight of the mouse liver (also see Supplementary Video S1). B. Patches demonstrate strong bioadhesion to the mouse liver. The detachment forces for the AuxES patches (also see Supplementary Video S2) are 10-fold higher than GelMA alone and robustly higher than those for a GelMA patch attached with an FDA-approved fibrin glue (TISEEL, Baxter). ***denotes p < 0.001 as analyzed by One-Way ANOVA followed by Tukey’s post-hoc tests. C. Schematic showing that the combination of hydrogen, covalent, and ionic bonding imparts instantaneous and strong bioadhesion and ultra-elasticity to AuxES patches.
AuxES patches are designed to adapt to the mechanics of dynamic organs
Dynamic organs have highly variable anisotropy and auxetic or non-auxetic characteristics. For example the lungs, myocardium, bladder, stomach, intestines, diaphragm, and the tendons are dynamic organs, of which the diaphragm and tendons are non-auxetic (positive Poisson’s ratio of surface deformation), while other organs are auxetic (negative Poisson’s ratio of surface deformation) (1). For the design of the AuxES patches, we considered intrinsic organ anisotropy, where the organ stiffness (E) is different in the longitudinal (assumed as the predominant direction of ECM fiber orientation) and transverse (direction perpendicular to the ECM fibers) directions. Since the void space between the ECM fibers reduces the resistance to deformation in the transverse direction (1), the longitudinal:transverse stiffness ratio (henceforth referred to as EL:ET) is greater than 1. Except for the lung, which is mostly isotropic (EL/ET ∼ 1), other organs feature anisotropy (EL/ET > 1) (1).
To determine the mechanical characteristics of AuxES patches, structural mechanics computational models were developed to simulate unidirectional (longitudinal or transversal) patch stretching. To establish the material properties of the materials used in the computational model, we first empirically measured the tensile strength of dog bone-shaped patches of different formation (Figure 3A). The 80% v/v GelMA – 20% v/v ACA combination demonstrated the highest stretchability (∼110% yield strain) and elastic modulus (∼20 kPa) and was therefore used in the computational models, which were then used to derive patch stiffness in the longitudinal and orthogonal directions sequentially. While many auxetic and non-auxetic architectures have been described in the literature (21–23), we chose the commonly described re-entrant (10, 24), sinusoidal ligaments (25), lozenge truss (26), rectangular truss (26) and honeycomb (27) architectures. Of note, these auxetic patch architectures have been explored our previous work(11), but we needed to re-implement the design schemes and the associated computational modeling for the new patch material to be able to derive the optimized patches for different dynamic organs. Here, we also explored some non-auxetic designs to be able to derive patches optimized for conformation to diaphragms and tendon tissue. We created a library of 64 designs demonstrating a broad range of auxetic, non-auxetic, and anisotropic properties by changing the width (w), height (h), radius (r), inclination angle (ia), and thickness (t), while keeping the overall patch size at 30 × 30 mm2 (Figure 3B, see Table S1 for the values of the parameters for each design iteration of the patches). Using this systematic approach, the patch stiffnesses and Poisson’s ratios were determined in the longitudinal and transverse directions (select computational outcomes are shown in Figure 3C), as summarized in Figure 3D, to map the different patch architectures to the dynamic loading of different organs. We also validated whether these patch architectures could withstand the different deformations of different organs by plotting the yield strains of these patches, Figure 3D, shows that the yield strains of the patch formulations for different organs were always in excess of the maximum organ deformation. This way, that the patches would be able to maintain their elasticity throughout repetitive cycles of organ deformation.
A. Different dog bone-shaped patch formulations first underwent tensile testing to optimize the patch formulation and its mechanical properties for computational modeling. 80% v/v GelMA – 20% v/v ACA demonstrated the highest stretchability (∼110% yield strain) and elastic modulus (∼ 20 kPa) and was used in the computational models. B. Selected auxetic (re-entrant honeycomb (10), sinusoidal ligament (25), and lozenge truss (26)) and non-auxetic (rectangular truss (26) and honeycomb (27)) architectures were used within the computational models. Design features (w, h, r, ia, t, te) were varied (see Table S1 for the values) to generate a library of over 60 patch designs. C. Computational outputs depicting internal stresses within the selected iterations of the geometries of auxetic (negative Poisson’s ratio) and non-auxetic (positive Poisson’s ratio) architectures using the 80% GelMA – 20% ACA formulation. Note that the auxetic patches expand laterally when stretched, while non-auxetic patches contract laterally when stretched. D. Plots of the patch stiffness ratios, Poisson’s ratios, and yield strains and overlap with the properties of individual organs. The yield strain plot lists the maximum linear strain that the organ undergoes in its normal physiological cycle. The patch fit to each dynamic organ was decided based on both the stiffness curves and Poisson’s ratios. The selected patch architectures are further verified such that yield strain of the patch is always greater than that of the normal physiological strain of the organ.
Collectively considering the stiffness and Poisson’s ratios of the different organs, and that the patch should exceed the maximum organ deformation, we finalized a patch architecture and its iterations for each organ (Figure 3D and Figure 4A). These patches were all ultra-stretchable (Figure 4B). We next performed tensile testing of the selected patch geometries for different organs to validate the computational predictions, which revealed a close correlation between the computational estimates and the experimental outcomes (Figure 4C). The yield strains of the patches were slightly higher experimentally than in silico, which might be because the computational models used perfect geometries with sharp edges, which can give rise to concentrated stress, whereas the sharp edges used experimentally had a fillet radius due to process limitations, which might reduce internal stresses.
A. Printed patches based on the selected patch designs for different dynamic organs. Note that some are auxetic and some are non-auxetic, depending on the target organ. B. Demonstration of patch ultra-elasticity (see Supplementary Video S3). C. Tensile testing of the selected patches for the different organs demonstrates a close correlation between the computational and experimental outcomes of the elastic modulus (stiffness) and the yield strains of the patches. D. Ex vivo implantation of the patches (15) over rat lungs, and comparison of a patch with square holes vs an auxetic patch (Loz itr5). Auxetic patches substantially increased their surface area by increasing the size of the voids, thereby conforming to the lung mechanics. E. Ex vivo implantation of the patches (30×30 mm2) over the right middle lobe of porcine lungs, where auxetic patches (Loz itr4) better conformed to the lung mechanics compared with non-auxetic (honeycomb) patches (see Supplementary Video S5). F. Patches were implanted in vivo over beating rat hearts (see Supplementary Video S6). The attached patches conform to the heart beat over several systolic and diastolic cycles. Lengths of patches across different cycles were measured via analysis of video frames in ImageJ. **represents p < 0.01 (group-wise comparisons by Tukey’s HSD post-hoc tests).
We next tested whether the patches could conform to a dynamic organ by determining the change in organ surface area, a direct correlate of the change in patch surface area. Since each organ has a distinct shape that further varies throughout the organ, we used volumetric fold change of the organ as the rational criterion for determining the change in organ surface area (Figure S1A). Of the different organs, the human stomach undergoes the highest dilation (up to 2.2-times its original volume) (1), so we used the stomach as the exemplar by developing a mechanized setup to inflate and deflate a stomach-mimicking balloon (Figure S1B). The air flow rates inside and outside of the balloon were finely calibrated such that the expansion of the balloon mimicked the surface area change of the stomach and attached the selected stomach-mimicking Re-en itr1 patches onto the balloon model and studying the patch mechanics. Upon increasing the balloon volume to approximately 2.2-times its original volume (45 to 100 cm3) using cyclical pressurization through air, we demonstrated that the auxetic Re-en itr1 patches easily conformed to balloon expansion and contraction (see Supplementary Video S4) and had a similar expansion ratio to that of the auxetic surface of the balloon. In contrast, a non-auxetic patch (Rect itr5, which is designed to mimic the mechanics of the muscular portion of the diaphragm (see Figure 4A) did not conform to the balloon expansion and demonstrated a significantly smaller expansion ratio. These experiments indicate the superiority of the auxetic patches over a non-auxetic patch, in applications concerning volumetrically deforming dynamic organs (bladder, stomach, heart, intestines, lungs). However, we would like to note that the balloon surface is not the same as serosal surface of the organs and this model is also unable to exhibit the intrinsic spatial and directional anisotropy of the organs (1). Therefore, to further ascertain the superiority of the patches, we used ex vivo lung models and in vivo heart models in the next study.
AuxES patches demonstrate conformation to different dynamic organs
An important criterion for any dynamic organ patch is the need to withstand the cyclical stretching and contraction and the resulting fatigue loading after administration. We therefore used ex vivo lung (rodent and porcine) and in vivo heart (rodent) models to test patch compliance. For the rodent models, we scaled-down the patches by a scaling factor of s = 0.5 (i.e., 15 × 15 mm2 patches with proportionally reduced features) and in porcine models we used a scaling factor s = 1.5 (i.e., 45×45 mm2 patches). Herein, the scaling of the patches will not affect the stiffness and Poisson’s ratios of the patches (1), since all the features are uniformly scaled, and their relative dimensions (e.g., the width and height ratios, etc.) remain the same. Unlike non-auxetic patches, the auxetic patches demonstrated a greater increase in overall surface area during physiological ventilation (PV) and hyperventilation (HV) states (Figure 4D, E, p < 0.05) in both the rodent and porcine models (also see Supplementary Video S5). Next, we applied the patches over the beating heart within an in vivo rodent model. The patches easily complied with the rapid motion of the heart and demonstrated rhythmic stretching and contraction in synchrony with the diastolic and systolic cycles of the heart (Figure 4F, also see Supplementary Video S6).
Drug-loaded patches designed with high flexibility and stretchability for wound healing
Compared with dynamic organs such as the lung and the heart, the skin undergoes complex deformations during normal physiology. For example, the sole and the dorsum undergo the highest deformation during dorsiflexion and plantarflexion, respectively, and the Poisson’s ratio is negative or positive across different regions (28). Such a large variation in skin mechanics mandates the use of region-specific auxetic or non-auxetic patch architectures. Figure 5A demonstrates the high variability in the Poisson’s ratios in different skin regions, in a synthetic anatomical foot model, from plantarflexion to dorsiflexion of the sole and the dorsum (28). In some regions, the Poisson’s ratio can be close to -0.6 (central region in the heel portion of the sole and toe portion of the dorsum), while other regions of the foot undergo minimal flexing. Thus, auxetic patches will be required to allow for flexible movement of the feet. Figure 5B demonstrates conformation tests of the patches in the toe regions of the sole (Re-en itr1 most suitable) and dorsum (Re-en itr4 most suitable) during dorsiflexion and plantarflexion in an anatomical foot model. The patches demonstrated excellent conformation to the movement of the foot (see Supplementary Video S7); for example, on the sole, Re-en itr1 expanded along both the length and width during dorsiflexion, and contraction along both the length and width during plantarflexion.
A. There is high variability in patch deformation under different physiological motions of the skin. Taking the sole and dorsum as example regions, the Poisson’s ratios of various regions during dorsiflexion or plantarflexion are indicated. The printed AuxES patches relevant to each sub-region are highlighted. B. Patch deformation relevant to the different subregions of the sole or dorsum in an anatomical foot model under dorsiflexion or plantarflexion. The patches can conform to the skin dynamics in each case (see Supplementary Video S7). C. Due to their intrinsic bioadhesiveness, AuxES patches allow the attachment of fluorescently labeled MSC-Exo to their surface. After 1 h of incubation in an exosome-rich solution, exosome attachment to the AuxES patches is substantially greater than that of pure GelMA patches. D. In vitro scratch assays with 3T3 fibroblasts demonstrate that patches laden with both CNPs and MSC-Exo clearly outperform patches laden with CNPs and the control groups (no patches). While the control groups only demonstrated ∼40% wound healing after 48 h (days 0 to 2), CNP-laden patches demonstrated up to 90% healing and the MSC-Exo and CNP-laden patches demonstrated complete wound healing.
After establishing the auxetic lattice design suitable for skin applications we next optimized drug loading onto the patches to promote wound healing. In our design, we specifically chose CNPs over other UV absorptive agents (such as FD&C yellow food dye (18)), as CNPs have added therapeutic benefits through their intrinsic anti-inflammatory (29), anti-microbial (30), anti-oxidant (29), and free-radical-scavenging (31) properties. We further hypothesized that we could exploit the intrinsic bioadhesive characteristic of the AuxES patches coated with biological therapeutics to further enhance their therapeutic potential. For this, we selected mesenchymal stem cell-derived exosomes (MSC-Exo) as the biological therapeutic (32– 34), since MSC-Exo have been shown to have intrinsic regenerative effects by inhibiting inflammation (35, 36), and promoting cell proliferation (37, 38), potentially making them an effective treatment for wound healing. We first assessed whether MSC-Exo could be successfully attached to the patches by dipping the patches into a solution containing fluorescently-labeled MSC-Exo (1011 MSC-Exo/ml), followed by quantification. Our AuxES patches with active bioadhesion clearly outperformed pure GelMA patches (Figure 5C). After 30 min in the exosome-rich solution, GelMA patches only fluoresced at the edges due to negligible bioadhesiveness, while AuxES patches demonstrated MSC-Exo coating across the entire patch surface.
We next used in vitro scratch assays (39) to evaluate the effectiveness of AuxES patches for facilitating wound healing. AuxES patches laden with both CNPs and MSC-Exo completely filled the scratch over 48 h, patches laden with only CNP resulted in ∼90% filling of the scratch, while control groups without any patch treatment only filled 40% of the scratches after 48 h (Figure 5D).
AuxES patches are effective at wound healing and promote wound healing response
The high conformation of AuxES patches to skin movements should prevent constrained movement after application and patch-induced mechanical stretch of the skin, lending themselves to the personalized treatment of burn wounds and diabetic ulcers with specific patches in different skin regions.
We next assessed the wound healing effectiveness of a week-long (i.e., day 0 to day 6) administration of the three groups of patches in a mouse model of cutaneous injury (Figure 6A). The patches attached easily to the mouse backs and conformed to the movement of the animal. After 24h of patch treatment, all groups containing AuxES patches (with or without CNPs or MSC-Exo) demonstrated significant wound healing compared with controls (no patch), which could be attributed to the enhanced cellular adhesion ligands (Arg-Gly-Asp (RGD) present in the GelMA which supports cellular proliferation (16)). After 3 to 4 days, AuxES patches laden with CNPs and MSC-Exo demonstrated significantly better wound healing (p < 0.05) compared with controls and animals with patches without CNPs, indicating that the CNPs and MSC-Exo also contribute to short-term wound healing. However, after a week, only AuxES patches with CNPs and MSC-Exo demonstrated significantly increased wound closure compared with other groups (p < 0.01). Hematoxylin and eosin staining of the periphery of the wounds indicated that the wounds treated with the AuxES patches with CNPs and MSC-Exo had improved re-epithelialization and increased granulation tissue in the wound bed (Figure 6B). Most samples, regardless of treatment group had a crust along the surface of the wound and adjacent epithelium which was comprised of proteinaceous fluid (interpreted to be serum) with large numbers of degenerating neutrophils (serocellular crust). In the wounds treated with AuxES patches with CNPs and MSC-Exo, there was also granulation tissue (fibrovascular tissue – normal component of wound healing) evident within the wound bed, which was minimally present or absent in the other groups. MSC-Exo have been shown to promote cutaneous wound healing through macrophage polarization to the anti-inflammatory M2 phenotype (40) and by promoting collagen secretion and angiogenesis (41, 42), which likely synergized with CNPs to provide robust wound healing in the AuxES patches containing both the therapeutics. Despite potential artifacts of sampling, it appeared that no other groups had evidence of notable re-epithelialization or granulation tissue formation. There was no evidence of a foreign body response to the AuxES patch material (Figure 6B).
A. AuxES patches laden with CNPs and MSC-Exo demonstrated significantly greater wound closure compared with other treatment groups after a week of administration over cutaneous wounds in mice. B. Photomicrographs of the margin of skin wound (black arrowhead denotes wound margin). Re-epithelialization (arrow) and granulation tissue (*) are only evident in wounds treated with the curcumin and MSC-Exo loaded AuxES patch. Hematoxylin and eosin (H&E) staining, 10x objective (scale bar = 100μm). *indicates p< 0.05, **indicates p < 0.01, ***indicates p < 0.001, and ****indicates p < 0.0001 (group-wise comparisons by Tukey’s HSD post-hoc tests). C. Heatmap depicting expression changes normalized to the no-patch treatment group, with expression levels equaling “No Patch” colored white. Upregulated genes are depicted in dark blue, down-regulated genes highlighted in yellow. D. Gene ontology (GO) biological process pathways expressed as fold enrichment as compared to mus musculus gene list. E. Zoomed in heatmap from panel (a) highlighting the genes annotated as mediators in the “proliferative” or “remodeling phases of wound healing. Identical color scale to larger heatmap. F. Reactome pathway analysis depicted with the -LOG of the entity’s ratio, or the ratio of genes provided vs the total pathway components. G. Volcano plot of differential gene expression, with the black points indicating the AuxES patch without CNPs patch compared to the no-patch group, and red as AuxES patch with both CNPs and MSC-Exo compared to no patch. The Y-axis is the negative log10 of the p-value as calculated with multiple t-tests for each gene. The X-axis is the log2 fold-change in expression of a given gene. ****p < 0.0001, ***p < 0.001, **p < 0.01, *p < 0.05.
Next, we sought to understand the biological response in the cutaneous wound as a result of the AuxES patch treatment. We employed Nanostring technology, which is an amplification-free technology that measures gene expression by counting mRNA molecules directly. Only genes that were at least 2-fold up or downregulated were considered biologically relevant. Pathway analyses on the genes in the AuxES patches containing only CNPs and with both CNPs and MSC-Exo showed that the patches were promoting a wound healing and immunological response. Because dermal wound healing occurs typically in phases, we stratified the genes into inflammation, proliferation, and remodeling (43) to assess the therapeutic effect of patch treatment (Figure 6C). The genes identified in this study were stratified via bioinformatic protein pathway software (Pantherdb and Reactome) as well as literature sources (44–51). As visualized in Figure 6C, the majority of the up and down regulated sequences fell into the “inflammation phase” of wound-healing. Of all the markers assessed in treated groups, only the markers from mice treated with the AuxES patches containing both CNPs and MSC-Exo showed greater than 3-fold upregulation (Figure 6G). The most significantly upregulated inflammatory proteins within the AuxES patches containing both CNPs and MSC-Exo were CCR4, ACKR4, CD7, IL-17b and IL-2. The majority of these genes have been implicated in promoting the transition from an inflammatory to a proliferative environment. For example, the upregulation of CCR4 (CD194, a receptor associated with chemotaxis, such as MCP-1 and RANTES) has been shown to promote wound healing through the recruitment of T-regulatory cells to the wound site, and is thus a precursor to the proliferation phase of wound healing (52).
ACKR4 has been found to be critical in promoting the egress of dendritic cells, as well as in the removal of inflammatory cytokines (53). The IL-17 family has been identified as key mediators of leukocyte infiltration and T-helper cell-mediated inflammation (54). Finally, within diabetes and lupus wound models, dermal wounds did not heal in IL-2 knockout models. Furthermore, where IL-2 was present, the skin regrowth was stronger, and T-cell expansion was more highly regulated (55). Furthermore, we observed that within the later wound-healing phases indicated as “proliferative” and “remodeling”, the AuxES patches containing both CNPs and MSC-Exo patches alone had nearly a third of the proteins that were upregulated, and this is visualized in (Figure 6F). Pathway analyses of these upregulated proteins show primarily immune-cell activation and interleukin signaling (Figure 6D,F).
After further analysis, we discovered that the AuxES patches were upregulating both proliferative and remodeling wound healing phase markers to a greater extent than any of the other treatments. This, in combination with the visual improvement in healing seen with the AuxES patches containing both CNPs and MSC-Exo, indicates that the patches are indeed promoting a more rapid wound healing response than the other treatments.
Void-filled AuxES patches are an effective treatment of puncture wounds
While auxetic patches are well suited to many dynamic organ pathologies (e.g., myocardial infarction treatment (10, 25)), they are unsuited for pathologies involving loss of bodily fluids (9) or air (pneumothorax (56)). We therefore modified the AuxES patches using a lower concentration GelMA-ACA mixture to create a first-of-its kind, hole-filling, auxetic patches by filling the voids in the auxetic lattice within the patches (Figure 7A). Using the same material to fill the voids can be a significantly simpler and cost-effective alternative to tissue glues (such as Fibrin) which are commonly used as sealants. Optimization of the concentration of void-filling (VF) material to successfully fill puncture wounds while still retaining the auxetic properties is shown in the Supplemental information (Table S2 and Figure S2). In this formulation, the 80% v/v GelMA – 20% v/v ACA mixture used for the patch lattices was further diluted to 50, 60, 70, 80, and 90% of its original concentration by mixing with PBS and application to the patch voids followed by UV exposure. Void-filling material diluted to 80% of its original concentration could allow the patch to retain its auxetic properties (see Supplementary Video S8 demonstration of the conformation of VF-AuxES patches to the deformation of a balloon; Figure 7B shows the balloon dilation measurements of the optimized VF-AuxES patches), while maintaining its integrity during expansion. This novel combination of void-filling materials into patches could be an effective strategy for the treatment of pulmonary air leakage (Figure 7C).
A. VF-AuxES patches were fabricated by filling the voids of AuxES patches using a reduced concentration of matrix containing GelMA and ACA. B. AuxES patches synthesized using a 20% lower concentration of material than that used for the lattices (80% v/v GelMA and 20% v/v ACA solutions) demonstrated auxetic properties on balloon deformation without opening matrix pores (also see Supplementary Video S8). C. Concept of using VF-AuxES patches for the treatment of pneumothorax in an in vivo rat model. D. Lung-mimicking VF-AuxES patch and its placement over rat lung (also see Supplementary Video S9). E. After inducing pulmonary injury, the tidal ventilation pressure during inspiration of the lung reduced from 7.2 to 5.8 mm H2O. VF-AuxES patches fully restored the ventilation pressure to physiological levels, while simple AuxES patches were unable to improve the lung ventilation due to air leak through the empty voids within the patch lattices.
We further explored the use of VF-AuxES patches (Loz itr4 geometry) for restricting pulmonary air leakage in an in vivo rat model. Pulmonary air leakage was induced in SD rats by open chest surgery and puncturing the lungs with an 18G needle. The instantaneous and strong bioadhesiveness of the patches allowed patch placement while the lung underwent normal physiological motion (Figure 7D, also see video of VF-AuxES patch over lung in Supplementary Video S9), a clear advantage over conventional patches requiring the lung to be static for application (57). We compared the change in tidal ventilation pressure in injured lungs without treatment or after treatment with simple AuxES patches (Loz itr4, without void-filling material) and VF-AuxES patches. We did not include a non-auxetic patch or a VF-non-auxetic patch as a control, as we had already established in the ex vivo rat models (Figure 4D) that the auxetic patches were superior to a non-auxetic patch for conformation to lung mechanics. Healthy lungs demonstrated a ventilation pressure of 7.2 mm H2O under dilation, which was reduced to 5.8 mm H2O in injured lungs (Figure 7E). As expected, there was no improvement in lung function in the simple AuxES patch group, but VF-AuxES patches fully restored lung function (Figure 7E).
Discussion
There has been a steady rise in the development and use of therapeutic patches (3, 58–60) over the last few decades for a wide range of applications including myocardial infarction (61), chronic wounds (62), organ hemorrhage (8, 9), and cancer (63, 64). However, most of these patches do not readily adhere to wet and dry tissues and do not consider the complex mechanics and volumetric changes of dynamic organs caused by their auxetic and anisotropic properties. Here, we present a novel AuxES patch composition and design that addresses the limitations of conventional patches. Our AuxES patches exhibit some unique properties which allow their easy adaptation towards dynamic organs: 1. ultra-elasticity (stretchable up to 400% its length without breaking); 2. instantaneous bioadhesion to dry and wet tissue (strong bonds form in < 1 s over wet tissues) with detachment forces are higher than patches applied via commercial fibrin glue); 4. cost-effective, off-the-shelf components which are compatible with photo-projection based printing; 5. wound healing capability in vivo (through the controlled delivery of CNPs and conjugated MSC-Exo); and 6. possible synthesis with void filling to heal puncture wounds.
The defined (375 μm) resolution derived through the optimized CNP concentration allowed the printing of patches that closely mimicked their intended designs, and closely recapitulated the simulated mechanical properties in tensile tests (Figure 4C). Resolution could be further improved by reducing the voxel size of the incident light beam (17). While the current system used chain-growth polymerization of the GelMA matrix in the presence of LAP, future research could use step-growth polymerization using thiol-ene photoclick chemistry (65, 66) to allow for both higher resolution fabrication and quicker fabrication times. One could also deploy higher resolution techniques such as volumetric printing (67, 68) or stereolithography (69) to achieve resolutions of up to 50 μm. For other biomedical applications, patches without CNPs can also be fabricated using food dyes as the absorptive agent. The patch combination composition is simple enough to easily replicate and fabricate and can be linearly scaled-up for large-scale patch fabrication. Since the patches can be manufactured in a single layer, custom masks can also be used with UV lamps (70) to fabricate the patches, negating the need for complex projection systems in 3D printers and further reducing the cost of fabrication.
Moreover, the AuxES patches are one of the few that instantly and strongly adhere to wet tissue, as shown in Figure 2 and Figure 4. These properties are important for future clinical translation of the technology, since organs do not need to be immobilized for several minutes for their application, a current requirement for many patches reliant on the formation of amide bonds for crosslinking (8, 9). While other mussel-inspired formulations relying on hydrogen bonding (71) and free radical polymerization (72) have demonstrated instantaneous adhesion, they rely on dopamine and its catecholic groups, which frequently undergo oxidation in neutral and basic conditions and require a fine redox balance to maintain tissue adhesion. This is difficult to achieve and severely compromises their adhesion and limit their practical applications in medicine (73). Further, our rational design framework for organ-specific soft anisotropic-auxetic patches ensures fatigue resistance over long-term administration while preventing undue stress on the organ.
Our wound healing experiments strongly suggest that a dual-therapeutic patch system containing CNPs and MSC-Exo is highly effective for promoting wound healing. Notably, AuxES patches demonstrated gradual dissolution into the skin after a week (Figure 6A) and does not need to be removed. A degradable patch is desirble as it would not need to be surgically removed after the wound has healed. However, a fast degradation may be less advantageous for deep-scar wounds or chronic injury such as myocardial infarcts (74). Here, the disintegration rate can be finetuned and controlled through photo-crosslinking density (14) and thus tailored to different types of therapeutic application. A higher density of covalent crosslinks within the patches could be used to prolong the time to dissolution and could be fine-tuned for different applications by increasing the degree of methacrylation and concentration of GelMA or the content of photoinitiator and UV exposure (14). The patch system containing CNPs and MSC-Exo was highly effective at promoting a wound-healing response, as shown by the studies in Figure 6. Dermal wounds go through a four-stage process of wound-healing; (1) hemostasis (∼0-12h after injury), (2) inflammation, (3) proliferation, and (4) remodeling. Because the injury model presented in this work extends to > 5 days, hemostasis is not under consideration as a major component of the observed wound healing. The inflammatory phase takes place anywhere from 1-8 days after injury, and is demarcated via expression of interferons, leukocyte migration, activation of M1 macrophages, natural killer cells, and dendritic cell infiltration (44–49). While chronic inflammation can lead to scarring and formation of fibrotic tissue because the time course of our injury models is less than ten days, the inflammatory markers observed are that of a normal wound healing process. The proliferation phase can begin as early as day 1, and lasts up to 8 days post injury (49–51). The treatments containing both CNPs and MSC-Exo had significantly more up-regulation of proliferative genes than the other two treatments, which we hypothesized is indicative of a more advanced stage of wound healing. Furthermore, the treatments CNPs and MSC-Exo also had a higher expression of remodeling phase markers (Figure 6E). The remodeling phase begins on day 3 and can extend for months after the injury, so an upregulation in these markers is indicative of a more advanced stage of wound healing (49–51).
Only a few patches have been clinically approved so far (75, 76), highlighting the challenges to clinical effectiveness and translation. Other adhesive materials have also been investigated with promising strong and instantaneous adhesion (77) and ultra-elasticity (78), but the materials lack rationally designed anisotropic and auxetic architecture, ultra-elasticity, instant bioadhesion, and compatibility with high resolution printing techniques. Regardless of the material used, our rational design framework demonstrated in this work can be used to design the anisotropic and auxetic patches to conform to specific organs, thus improving long-term integration of patches within hosts. With our VF-AuxES patches, we have demonstrated how this rational design framework can be further expanded to also include the treatment of dynamic wounds such as pulmonary air leakage. To our best knowledge, this is the first to demonstration of this combined scheme, which allowed us to retain the auxetic properties whilst also preventing gas leakage from the lung.
In conclusion, our patch technology provides a versatile platform for easy and cost-effective adaptation towards different dynamic organs, with potential applications ranging across broad range of pathologies including battlefield injuries, myocardial infarction, chronic wounds, pneumothorax and other diseases.
Materials and Methods
Patch materials synthesis
GelMA was prepared using existing protocols for the controlled methacrylation of gelatin (79). Briefly, 5 g of gelatin (Bloom 300, porcine derived, Sigma Aldrich, St. Louis, MO) was dissolved at 10% w/v in 0.25 M carbonate-bicarbonate buffer (79) containing 2% w/v of sodium bicarbonate and 0.12% w/v of sodium carbonate in deionized water. The gelatin solution was kept at 50°C until achieving a clear solution. Next, 159 μl of methacrylic anhydride (MA) was added and the reaction allowed to run for 60 min at 50°C. The reaction was stopped, and excess MA removed, by the addition 100 ml of 1:1 mix of pure ethanol and acetone. The degree of methacrylation was 40%, as determined by 1H NMR (79). Acrylic acid was used as purchased (79-10-7, Sigma Aldrich). CNPs were prepared using established methods (80, 81). Briefly, curcumin (Cur) powder (C1386, Millipore Sigma) was dissolved in tetrahydrofuran (THF) solution at 25 mg/mL. 50 μL of the Cur-THF solution was rapidly injected into 450 μL of deionized water with vigorous stirring at 1400 rpm to aggregate as nanoparticles. The CNP suspension was air-dried to remove the organic solvent and lyophilized. The resultant CNPs were stored at -20 °C until further use. To prepare the material for the patches, 10% w/v GelMA in PBS at 37°C and acrylic acid (79-10-7, Millipore Sigma) were mixed at different ratios (80%-20% or 90%-10%, also see results in Figure 3A), followed by the addition of 0.03% w/v LAP photoinitiator and the desired amount of CNPs (1 – 2.5 mg/ml) to formulate the photoink.
Patch printing and postprocessing
The material was cooled to room temperature (24°C), and 750 μl was added to a rectangular enclosure (50 × 50 mm2) attached onto a PDMS-coated petri dish and uniformly smeared across the enclosure using a pipette tip. Acrylic acid prevented the thermo-reversible crosslinking of GelMA, which is otherwise a native property of GelMA hydrogels (16). The petri dish was then placed above a digital light projection system (LumenX, Cellink AB) and projection printed using 405 nm UV light at 20 mW/cm2 for 2 min. Patches were then neutralized and further crosslinked by dipping into a solution containing 0.5M CaCl2 and 0.05M NaOH in deionized (DI) water. Before any further experiments, the patches were gently washed with DI water.
Tensile testing
The tensile properties of the patches were determined using an in-house setup consisting of a linear stage actuator (101-80-124, SainSmart) as the stretching mechanism, and a 5 kg load cell (TAL220B, Sparkfun) and amplifier (HX711, Sparkfun) as the stationary anchor. The linear actuator was controlled via a stepper driver (101-60-197, SainSmart) and motion controller (101-60-199, SainSmart). The ends of the patch were clamped between the linear actuator and the load cell, and the patches stretched 0.125 mm/s similar to previous existing studies (82, 83). The load cell was connected to an Arduino Uno and provided the stress vs. strain curve to determine the mechanical properties.
Bioadhesion measurements
Patch bioadhesion was measured over ex vivo mouse livers. Briefly, the mouse livers were gently dabbed using a blotting paper to remove excess blood followed by placement over the stationary load cell of the tensile testing apparatus. Next, patches were placed on the mouse liver. For GelMA patches administered using fibrin glue, the fibrin sealant was first prepared as per the manufacturer’s guidelines (TISSEEL, Baxter). GelMA patches were placed on the mouse livers, followed by placement of 50 μl each of the fibrinogen prepolymer solution and the thrombin crosslinker solutions to attach the patches to the mouse liver. All patches were administered such that one end of the patches hanged over the edge of the liver by 5 mm, which was attached to the linear translation stage of the tensile testing apparatus. The patches were stretched at 0.125 mm/s and the highest force generated during the stretching procedure was noted as the detachment force.
Evaluation of patch biocompatibility in vitro
NIH 3T3 fibroblasts (CRL-1658, ATCC) were cultured in Dulbecco’s modified Eagle’s medium (DMEM) and 10% fetal bovine serum and 1% penicillin-streptomycin in six-well plates (2 ml of medium per well) until 40% confluency. The patches were then added into the wells and the viability assessed after 2 days using the Live/Dead™ assay (L3224, ThermoFisher Scientific).
Patch designs
Solidworks (Dassault Systems) was used to prepare the patch designs. The inset images (Figure 3B) constituted the repetitive elements in the linear array of the lattice. The dimensional parameters (h, w, r, t, te, and ia) were defined as global variables, so that they could be easily altered to change the overall patch design. The overall size of the patches, irrespective of the dimensionality of the lattice elements, was kept at 30 × 30 × 1 mm3.
Computational modeling of the patches
The 3D patch geometry was imported into the structural mechanics module of COMSOL Multiphysics (COMSOL Inc.). A fixed boundary condition was applied on one edge of the patch (either the longitudinal (L, assumed along x-axis) or transverse (T, assumed along y-axis) direction) and a linear displacement of δ = 10% strain applied on the opposite edge. A mesh density of 1/10th of the minimum element size (0.25 mm) was selected as previously (84). The stiffness of the patch (E) was determined from the strain energy (US), strain (δL or δT= 10%), and volume of the patch (VP) as per E = δ2VP/2US. For some patches, this stiffness was different when stretched longitudinally (along x-axis), than when stretched transversally (along y-axis), thereby making the stiffness ratio smaller or larger than 1. The Poisson’s ratio (νp) was determined from the deformation transverse deformation when stretched longitudinally, hence νp = δT/δL. The yield strain (δYield) of the material was calculated at the stretch when the maximum internal stress (von mises = σi) within the patches exceeded the yield strength (σmax) of the bulk material: δYield = σi/σmax.
In-house balloon model to demonstrate patch compliance to organ mechanics
The balloon model consisted of a programmable control of a pneumatic valve using a linear actuator capable. An air flow meter was used to calibrate the air flow (VFA-26-BV, Dwyer Instruments). An open cylindrical tubing attached to the inlet of the balloon allowed the input air to exit the balloon. To measure the change in the surface area of the patches compared with that of the balloon, point marks were placed on the portion of the balloon encompassing the patches and around the patches. The relative expansion of the patch and the balloon was measured by calculating the distance between the opposite edges of the patches, and the opposite points placed along the balloon, in ImageJ. The product of the measured length (L) and width (W) was the surface area. The expansion ratio was the ratio between the change in surface area of the patches (LP×WP) to the change in surface area of the square encapsulated by the points on the balloon (LB×WB).
Ex vivo evaluation of patch compliance
Rat lungs were excised from anesthetized male Sprague Dawley (SD) rats (∼300 g average weight, 12 weeks old) after euthanasia. Porcine lungs (cold flushed post euthanasia) were obtained from North Carolina State University School of Veterinary Medicine (UNC Institutional Animal Care and Use Committee (IACUC) Protocol No: 20-045.0). Patches were attached to the lungs using forceps. The lungs were then ventilated using a rodent ventilator and a Servo-i mechanical ventilator, respectively. The ventilation volumes were 7.5 ml/kg (physiological ventilation) and 12.5 ml/kg (hyperventilation) for rat lungs, and patch motion was recorded as a video. The patch size was determined as for the balloon model.
In vivo evaluation of patch compliance
An intraperitoneal injection of ketamine/xylazine was used to anesthetize the SD rats and the rats (UNC IACUC Protocol No: 20-045.0) were kept unconscious while ventilating at a volume of 7.5ml/kg (typical for rodent studies (85)) with isoflurane and oxygen (after tracheotomy). The lungs and the heart were exposed via sternotomy. To demonstrate the compatibility of the patches with minimally invasive surgery, the patches were wrapped around a minimally invasive surgical scope and attached onto the heart. Patch motion was recorded as a video. The changes in patch length were determined using analysis of different frames of the captured video in ImageJ.
In vitro healing evaluation using scratch assays
Methods based on previous work (39) were used for the scratch assay. Briefly, 3T3 cell cultures within six-well plates were allowed to reach 100% confluency and were then scraped using a 23G needle to create ∼650 μm (wide) and 5 mm (long) scratches through the cells. The cells were then allowed to proliferate through the scratch without any treatment (control), or in the presence of patches (placed within the wells after creation of the scratch) containing only CNPs or both CNPs and MSC-Exo. The width of the scratch was determined after 24 h and 48 h using analysis of brightfield images in ImageJ.
In Vivo Murine Wound Healing Model
Female, 6-8 week old C57BL/6J were purchased from Jackson Laboratory. The animal studies were approved and carried out in compliance with the Institutional Animal Care and Use Committee standards. The mice were housed individually with 12 h light−dark cycles. For wounding, mice were anesthetized using gaseous isoflurane and received a subcutaneous injection of 0.05 mg/kg buprenorphine. Hair was removed from the dorsal region of the mouse using clippers and a depilatory cream, and the skin was prepared for surgery using betadine and 70% ethanol. A sterile 6 mm biopsy punch was used to outline a circular pattern between the shoulders. Forceps were used to lift the skin, and surgical scissors were used to create a full thickness wound on each mouse. Mice were randomly assigned into four treatment groups of 5 mice each and treated with AuxES patches without CNPs, AuxES patches with CNPs, AuxES patches containing both CNPs and MSC-Exo, or no patch. Patches were adhered to the wound site immediately following wounding. After the initial treatment, wounds were covered with a Band-Aid to facilitate patch adherence and prevent patch removal by the mice in the first 24 hours post-wounding and treatment. Each day, all wounds were measured in perpendicular directions using calipers for wound area calculations, and wounds were imaged. On day 6 post-wounding, mice were sacrificed, and residual wounds were harvested and stored at −80 °C.
Histopathology of in vivo skin wounds
At Day 6 after wounding, the wound and surrounding skin were collected from 3 mice per group. Samples were fixed in 10% neutral buffered formalin for ∼72 hours and routinely processed to paraffin, embedded, and sectioned to 5 μm, stained with hematoxylin and eosin, and evaluated microscopically. Wounds were trimmed at approximately the middle of the wound bed.
Pathway analyses of NanoString RNA sequencing data - immune gene detection and quantification
Isolation of mRNA was performed as previously described (86). Briefly, cells were lysed with TRIZOL buffer (Sigma) and total RNA was isolated by chloroform extraction and quantified using a nanodrop 2000™ spectrophotometer. NanoString technology and the nCounter Mouse Immunology Panel was used to simultaneously evaluate 561 mRNAs in each sample (87). Each sample was run in triplicate. Briefly, a total of 100 ng mRNA was hybridized to report-capture probe pairs (CodeSets) at 65°C for 18 hours. After this solution-phase hybridization, the nCounter Prep Station was used to remove excess probe, align the probe/target complexes, and immobilize these complexes in the nCounter cartridge. The nCounter cartridge was then placed in a digital analyzer for image acquisition and data processing. Hundreds of thousands of color codes designating mRNA targets of interest were directly imaged on the surface of the cartridge. The expression level of each gene was measured by counting the number of times the color-coded barcode for that gene was detected, and the barcode counts tabulated. nSolver v4.0, an integrated analysis platform was used to generate appropriate data normalization as well as fold-changes, resulting ratios and differential expression. Replicates (n=3) of sequence reads were normalized to the control (no patch) for downstream analysis. Sequences that were of identical expression levels in all samples were excluded from pathway analysis as background. Sequences that were calculated as being 2-fold higher or lower than the AuxES patch containing only CNPs, or AuxES patch containing both CNPs and MSC-Exo, were used for pathway analyses. These fold-changes in expression were analyzed using GraphPad Prism and statistics conducted using a two-way ANOVA for grouped analyses. Up- or downregulated sequences were analyzed using the Panther protein database overrepresentation test, under the GO-Biological process complete analysis, with Fisher’s Exact test and corrections via calculation of false-discovery rate.
MSC-derived exosomes (MSC-Exo) synthesis, fluorescent labeling, and attachment to the patches
The serum for the exosome-free medium was prepared by ultracentrifugation of fetal bovine serum (FBS, MilliporeSigma) at 100,000 x G for 12 hours at 4°C, followed by extracting the supernatant. This serum was then mixed at 10% concentration in Dulbecco’s modified Eagle Medium (DMEM, Millipore Sigma). Human bone-marrow MSCs (passage 2) were first cultured in 10% FBS and 90% DMEM until reaching 70% confluency in T-75 flasks. Then, the cells were washed three times with PBS to remove excess serum, and the medium was replaced with exosome-free media. The medium containing MSC-Exo was collected after 48 h and the MSC-Exo concentrated via differential ultracentrifugation (88), resuspended in PBS, and quantified by nanoparticle tracking analysis (NTA). To fluorescently label the MSC-Exo, an NHS ester fluorophore (Dylight650, ThermoFisher Scientific) was added at a concentration of 0.1 mg/1011 MSC-Exo, followed by incubation at 37°C. The excess fluorophore was then removed using exosome spin columns (molecular weight cutoff 3 kDa, ThermoFisher). MSC-Exo were then reconstituted in PBS at a concentration of 1011 exosomes/ml. On each patch (air dried for 5 min), 100 μl of the fluorescent-exosomes suspension was added to allow the exosomes to conjugate for 30 min at 24°C. The patches were then washed three times with PBS to remove unattached MSC-Exo. Patches were then cut into 4 sections and dipped in 24-well plates containing 250 μl of PBS, and the absorbance was measured at 647 nm using a plate reader. For confocal imaging (Fluoview, Olympus), the 647 nm laser was used to visualize the patch samples in PBS.
In vivo lung puncture wound treatment with composite AuxES patches
SD rats were prepared as for the in vivo compliance studies (physiological ventilation volume of 7.5 ml/kg was used, IACUC Protocol No: 20-045.0). The tidal ventilation pressure in the airway was continuously monitored with a pressure transducer (Transpac IV, ICU Medical Inc). Pulmonary air leakage was induced in the right lower lobe of the lung by inserting an 18G needle up to 1 cm deep in the lung tissue. Patches were placed over the injury site with forceps. Any changes in ventilation pressure were noted throughout the procedure.
Statistical analysis
All experimental data are presented as mean ± standard deviation. For in vitro, ex vivo, and live animal models, a sample size of n = 3 was used. Statistical tests were performed in GraphPad PRISM (GraphPad Software, La Jolla, CA). One-way ANOVA was used for the bioadhesion measurements (Figure 2C), balloon experiments (Figure S1B), loading of MSC-Exo over AuxES patches (Figure 5C), and pneumothorax treatment experiments (Figure 7F). Two-way ANOVA was used to assess patch conformation to ex vivo lungs (Figure 4D,E), and in vitro and in vivo wound healing studies (Figure 6C,D). Tukey’s HSD was used for the post-hoc analysis. In the Nanostring analysis, statistical analysis of the fold-change values of the different genes (see Supplementary Table S3) was conducted using GraphPad Prism v 9.3.1. Means of the no-patch fold-change were compared to the means of the Patch containing CNPs and MSC-Exo using multiple unpaired T-tests (one per row). Variance assumptions assumed individual variances for each row, and multiple comparisons were tabulated using False Discovery Rate (FDR), and the two-stage step up method (Benjamini, Krieger, and Yekutieli) with desired Q of 1.00%.
Acknowledgements
This work was in part funded by the Eshelman Institute for Innovation at the UNC Eshelman School of Pharmacy and the National Institutes of Health (R01CA241679, R01EB023262, and R21GM135853). Natalie Jasiewicz is funded through the PhRMA foundation predoctoral fellowship. Tom Egan and John Blackwell were supported by the UNC Lung Transplant Research Fund, with generous contributions by John Doherty and the Ferguson family, and the Cornelia D. Condon Memorial Fund for Lung Transplant Research. Some of the figures were created using Biorender.com. A patent is pending on the disclosed subject matter (US 63/326,982).
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