Tunable regional targeting of self-assembling peptide nanomaterials in acute myocardial infarction

Abstract

Acute myocardial infarction (AMI) causes cardiac tissue damage and creates a transient leaky vasculature, which is an attractive target for intravascular therapeutics. Nanomaterials have the potential to deliver therapeutics to the injured heart, yet the influence of particle size, shape, and surface decoration on infarct localization has not been systematically examined. Here, we evaluated a self-assembling peptide platform that can be tuned to form high aspect ratio nanofibers or globular nanoparticles by modifying the β sheet forming domain. This platform has tunable nanofiber length distribution via sonication and aging as well as surface decoration via targeting peptides. In a rat model of AMI, the incorporation of a type IV collagen targeting peptides significantly increased localization of the nanofibers to the injured leaky vasculature while globular peptides diffuse through and targeted the infarcted myocardium. These results demonstrate that this modular peptide nanomaterial can be tuned via geometry and surface moieties to either target to the compromised vasculature components or infiltrate the infarct. This study establishes size, shape, and surface decoration as critical features for effective intravascular nanomaterial delivery after AMI.

1. Introduction

Acute myocardial infarction (MI) is characterized by cardiomyocyte death, immune cell infiltration and inflammation^1-3. The resulting injury causes leaky vasculature that serves as a critical therapeutic target and entry point for intravascular therapeutics. Current treatments include reperfusion and various pharmaceuticals, which are limited by poor infarct retention and off-target effects⁴. Delivering therapeutics intravascularly with increased targeting to the infarct remains a significant challenge and opportunity for post-MI treatment.

Nanomaterials are advantageous due to their modularity, where properties such as size, shape⁵, targeting method³, surface chemistry⁶ and composition can be tuned to match a therapeutic need. Numerous targeting strategies have been explored, ranging from targeting cell-surface receptors to extracellular-matrix (ECM) components and inflammation-specific markers³. A hallmark of many inflammatory diseases, including MI, is vascular dysfunction⁷ which consists of endothelial activation, increased permeability, and exposure of the basement membrane. Exploiting these targets requires an understanding of how these nanomaterial properties impact localization and retention to the permeable vasculature⁸.

Among the tunable parameters, shape and size significantly impact vascular margination^7-10, where particles are displaced toward the vessel periphery through collisions with red blood cells and a variety of fluid flow effects. Larger nanomaterials^11,12 and anisotropic nanomaterials^13-16 marginate more efficiently than small, spherical counterparts, therefore increasing the local concentration at the endothelial surface and enhancing binding or extravasation to the injury site. Margination has been leveraged to target the vasculature in healthy animals¹⁷, and the inflamed vasculature in injury models like atherosclerosis¹⁸ and stroke¹⁹ but has not been systematically examined post-MI.

Our prior work demonstrated that an intravascularly deliverable decellularized extracellular-matrix (iECM) biomaterial targets the leaky vasculature generated by MI²⁰. The high-molecular-weight nanofiber component of iECM bound more efficiently both in vitro to basal-lamina ECM and in vivo to gaps in the endothelium than its globular peptide fraction⁹, suggesting that the anisotropic geometry combined with intrinsic ECM-binding drives this unique targeting. Inspired by these findings, we sought to recapitulate the shape dependent targeting capabilities seen in iECM nanofibers and selected a β-sheet-forming self-assembling peptide²¹ (ABA tri-block) that forms anisotropic nanofibers whose length distribution can be tuned with probe sonication and solution aging. The peptide backbone allows for facile incorporation of targeting peptides and we therefore added a type IV-collagen-binding peptide (T4C) ²² to target the exposed basement membrane of leaky vasculature. Using a rat model of acute MI, we systematically varied nanofiber length polydispersity and compared fiber versus globular structures to determine how size, shape, and surface decoration impact regional infarct targeting and retention.

Our study provides a scalable, peptide nanomaterial whose shape, size and surface decoration are tunable. Additionally, we quantitatively evaluate how nanofiber length distribution influences post-MI vascular targeting. Altogether, we establish a precedent for engineering intravascular nanomaterials that can be tuned either to bind the damaged endothelium or to diffuse into the infarcted myocardium. This work advances the fundamental understanding of how intravascular nanomaterial design influences localization for AMI therapy, particularly on how tunable biophysical and biochemical properties affect the targeting performance of such nanomaterials.

2. Results and discussion

2.1. Fabrication optimization of self-assembling peptide nanofibers

The previously developed self-assembling peptide – KKQLQLQLQLQLQLKK – (KQL6)²¹ was used to generate the nanofibers for this study. All peptides used are outlined in Table 1. Using probe sonication, we subjected freshly dissolved KQL6 peptide to increasing amounts of sonication time and sonication amplitude. These various samples were then used immediately for negatively stained transmission electron microscopy (nsTEM) imaging and quantitative analysis of nanofiber length. The nanofiber length distribution and average, calculated with the average nanofiber length per TEM image, became narrower (Figure S1a-b). Increasing sonication amplitude by 10% for the same amount of sonication time negligibly changed the nanofiber length distribution. The increase in sonication time from 2 minutes to 10 minutes significantly narrowed the length distribution; however, we discovered that this resulted in significant evaporation of the sample solvent and sample heating. Therefore, we opted to submerge our nanofiber samples vials in wet ice during the sonication process to reduce sample heating and solvent loss.

Table 1.

List of peptides, sequences from N to C terminus, and nomenclature.

Peptide	Sequence	Nomenclature
Self-assembling peptide nanofiber sequence	KKQLQLQLQLQLQLKK	Untargeted nanofiber
Type IV collagen targeting peptide nanofiber sequence	KLWVLPKGGGKKQLQLQLQLQLQLKK	T4C-targeted nanofiber
Type IV collagen targeting scramble peptide nanofiber sequence	PWKKLLVGGGKKQLQLQLQLQLQLKK	T4C-scramble nanofiber
Type IV collagen self-assembling peptide globular control	KLWVLPKGGGKKQLQLQLKK	T4C-targted globular
Type IV collagen scramble globular control	PWKKLLVGGGKKQLQLQLKK.	T4C-scramble globular

We next investigated the dependence of nanofiber length distribution on the solvent and targeting peptide, in addition to sonication and sample age. Previously, it was shown that when KQL6 is dissolved in 10 mM Tris buffer, the nanofiber reached an average length of 120 nm after 3 weeks of aging at room temperature²¹. Ionic strength and pH of the solvent can also be used to dictate the nanofiber length²¹. Since the nanofiber was intended to be intravascularly delivered, we opted with saline to match the salt concentration within the blood. Additionally, for the peptide to be fluorophore tagged for in vivo investigation without significant modifications to the sequence, the use of NHS-ester and primary amine chemistry was most convenient. With an NHS-ester chemistry, Tris buffer would competitively react with the fluorophore, reducing conjugation efficiency. We also explored incorporation of a targeting moiety. At areas of inflammation and ischemia, endothelial cell junctions are disrupted, leading to leaky vasculature and exposure of the basement membrane, which contains type IV collagen and laminins²³. We therefore hypothesized that the inclusion of a previously developed type IV collagen (T4C) targeting peptide²² would lead to binding in the gaps between endothelial cells in the leaky endothelium similar to the iECM²⁰. This targeting peptide was also selected due to its net positive charge, which would not interfere with the KQL6 self-assembly and backbone charge. The KQL6 peptide was used in hydrogel applications at 10 mg/mL or 5 mM peptide concentration^21,24. Since our intended delivery was intravascular, we investigated the nanofibers at 10- and 100-fold dilutions (500 and 50 μM or 1 and 0.1 mg/mL respectively) to reduce the risk for aggregation upon delivery. The distributions of the nanofibers without sonication at both 50 and 500 μM were similar between the untargeted nanofiber and T4C-targeted nanofiber (Figure S2a) (untargeted 50 μM: 36 ± 21 nm, untargeted 500 μM: 38 ± 27 nm, T4C-targeted 50 μM: 42 ± 27 nm, T4C-targeted 500 μM: 39 ± 21 nm). The average nanofiber lengths were also not significantly different (Figure S2b). These untargeted and T4C-targeted nanofibers were then subject to sonication and aging at room temperature (Figure S3a-b, S4a-b). At 50 μM, we found that both untargeted and T4C-targeted nanofibers reached a length distribution and average length plateau at 2 weeks (Figure S3a-b), which was earlier than previous findings of nanofiber aging²¹, highlighting differential solvent effects. The average nanofiber length was 139 ± 146 nm for 50 μM untargeted nanofibers and 127 ± 89 nm for 50 μM T4C-targeted nanofibers. We then followed the 500 μM untargeted and T4C-targeted nanofibers out to 2 weeks which resulted in similar average nanofibers at 146 ± 88 nm for 500 μM untargeted and 130 ± 61 nm for 500 μM T4C-targeted nanofibers (Figure S4a-b). The nanofiber length distributions between sonicated non-aged groups were similar, regardless of concentration and targeting peptides. For the 2-week aged groups (Figure 1a), similar distributions were seen as well. Across nanofibers compared to the same timepoint of aging, there were no significant differences (Figure 1b). When the sonicated non-aged samples were compared to the sonicated 2-week aged samples, there is a significant difference in both the distributions and the average nanofiber lengths (Figure 1a-b).

Figure 1. Nanofiber fabrication optimization.

A comparison of the a) distribution and b) average nanofiber lengths of untargeted and T4C-targeting nanofibers after sonication with 2 weeks of aging and without aging at different concentrations. For all statistics, one-way ANOVA with Tukey’s post hoc; ns p>0.05, *p < 0.05, **p < 0.01, ***p < 0.001, ****p < 0.0001. Sample size n= 5-8 images and n>400 nanofibers measured total. Mean ± standard deviation is plotted.

Since we sought to investigate the impact of nanofiber average size and polydispersity on the in vivo biodistribution post-MI, we classified three groups of nanofibers with varying processing parameters – 1) sonication and no aging, 2) no sonication and no aging, and 3) sonication with 2 weeks of aging – into 3 groups: 1) short, 2) medium and 3) long nanofibers respectively. The morphologies were investigated via TEM (Figure 2a-f) and length distributions (Figure 2g), and averages (Figure S5) were quantified. Within the untargeted nanofibers, the average length distribution for the short nanofibers were 23 ± 15 nm, medium was 38 nm ± 27 nm, and long nanofibers were 146 ± 88 nm. For the T4C-targeted nanofibers, the average length distribution for the short nanofibers were 29 ± 11 nm, medium was 42 ± 27 nm and long were 130 ± 61 nm. While the average values per TEM grid (Figure S5) were not significantly different, the distribution between the short and medium nanofibers were significantly different. The short nanofibers not only had lower averages, but also a smaller standard deviation showing how sonication can significantly reduce the polydispersity of the nanofibers. This allowed for the comparison of not only the length dependence, but also the impact of polydispersity. Moreover, the medium nanofiber constitutes the most straightforward, translatable biomaterial platform, because it requires no additional processing after dissolution in a suitable solvent as compared to the short and long nanofibers.

Figure 2. Nanofiber length distributions at different orders of magnitude. TEM micrographs of nanofibers with and without sonication and aging.

Sonicated non-aged short nanofibers that are a) untargeted and d) type IV collagen (T4C) targeted. Freshly resuspended, non-sonicated medium nanofibers that are b) untargeted and e) T4C targeted. Sonicated and 2-week aged long nanofibers that are c) untargeted and f) T4C targeted. Quantified g) distribution of the short, medium and long nanofibers. For statistics, one-way ANOVA with Tukey’s post hoc; ns p >0.05, *p < 0.05, **p < 0.01, ***p < 0.001, ****p < 0.0001. Sample size n= 5-8 images and n>400 nanofibers measured total.

These data confirm previous reports that the KQL6 self-assembling nanofiber platform can successfully incorporate targeting peptides^24,25, as seen through our addition of the T4C targeting peptide to the KQL6 backbone. The use of saline over Tris resulted in higher average nanofiber lengths and faster nanofiber aging equilibrium further supporting previous results of solution ionic strength effects on the nanofiber equilibrium²¹. Finally, the use of sonication and aging can be used to tune the nanofiber length average and polydispersity of the nanofiber sample. Altogether, we show the tunability of this self-assembling peptide nanofiber platform.

2.2. Length dependent localization of self-assembling peptide nanofibers

Prior to in vivo use, the nanofibers were subject to cytocompatibility experiments with L929 murine fibroblasts (Figure S6a-b). Proliferation of the fibroblasts resulted in a noticeable decrease in activity above 50 μM of peptide (Figure S6a). Therefore, the maximum in vitro tolerable does of the peptide was 50 μM. Due to the approximate 10-fold dilution in concentration of intravascularly delivered nanomaterials within the blood volume of rats, an injection concentration of 500 μM was selected to maximize the delivered dosage since the total blood concentration after injection would be approximately 50 μM and therefore at the maximum tolerable dose found in vitro.

The self-assembling peptides were evaluated for osmolarity, colloidal oncotic pressure and hemolysis (Figure S7a-c). The osmolarity of the surrounding fluid determines the likelihood of hemolysis and the colloidal oncotic pressure is the osmotic pressure exerted by large molecules, primarily plasma proteins, and is responsible for holding water within the vascular space^26,27. Typical osmolarity of physiological, isotonic conditions is around 275-295 mOsm/kg of water²⁶ while human plasma colloidal oncotic pressure is approximately 20-25 mmHg²⁷. All the self-assembling peptides exhibited average osmolarities between 310-340 mOsm and all the colloidal oncotic pressures were more than double that of saline suggesting a hypertonic state and a potential for hemolysis (Figure S7a-b). For the hemolysis percentage (Figure S7c), while all the self-assembling peptides showed a significant increase in hemolysis relative to saline, all the values were less than 2%, which based on current ISO standards, classifies this biomaterial as nonhemolytic (0-2%)²⁸.

In a rat MI model, the short, medium, and long untargeted and T4C-targeted nanofibers and saline were delivered via tail vein injection upon reperfusion. This delivery protocol simulated therapeutic delivery shortly after a balloon angioplasty reperfusion procedure commonly done in patients²⁹. At 1 hour after tail vein injection, the hearts (Figure 3) and major satellite organs (Figure 4) were harvested, fluorescently scanned, and quantified to investigate biodistribution. The short T4C-targeted nanofiber was retained within the left ventricle with a significantly higher signal compared to saline. The quantified average signal intensity (Figure 3a) showed a decrease in T4C-targeted nanofiber retention with increasing T4C-targeted nanofiber length suggesting a length-dependent relationship. The untargeted nanofiber did not show any length-dependent differences. When the left ventricle to right ventricle ratio (LV/RV) was compared, only the short T4C-target nanofiber had a significantly higher infarct retention over saline although numerical averages for other targeted groups were also higher (Figure 3b).

Figure 3. Impact of nanofiber size on self-assembling peptide left ventricle retention 1 hour after injection.

a) The left ventricle retention of the untargeted and type IV collagen targeting nanofiber at three different length distributions. b) The quantified fluorescent retention of the nanofibers within the left ventricle. c) The left ventricle (LV) to right ventricle (RV) ratio of signal for the various groups. For all statistics, one-way ANOVA with Tukey’s post hoc. ns p>0.05, *p < 0.05, **p < 0.01, ***p < 0.001, ****p < 0.0001. Sample size n= 3-4 biological replicates. Mean ± standard deviation is plotted.

Figure 4. Impact of nanofiber size on self-assembling peptide satellite accumulation 1 hour after injection.

a) The biodistribution of the untargeted and type IV collagen targeting nanofiber at three different length distributions. b) The quantified fluorescent retention of the nanofibers within the satellite organs. For statistics, two-way ANOVA with Tukey’s post hoc ns > 0.05 not shown, * p < 0.05, ** p < 0.01, *** p < 0.001, **** p < 0.0001 with comparison of row effects between treatments of only the same organ. Sample size n= 3-4 biological replicates. Mean ± standard deviation is plotted.

Quantification of the satellite organs showed that the untargeted nanofiber was predominantly retained within the lungs compared to the other organs for all lengths (Figure S8). Previous work showed that multivalent anions can crosslink the untargeted nanofiber into hydrogels^30,31. Since the nanofibers must pass the lungs through the capillaries within the alveoli prior to delivery to the rest of the body, the combination of phosphate driven crosslinking and reduced flow rate within capillaries may have caused accumulation of the untargeted nanofibers in this low flow region therefore reducing infarct retention. The T4C-targeted nanofibers showed lung retention in a length dependent manner; longer T4C-targeted nanofibers were significantly retained within the lungs while the shorter T4C-targeted nanofibers likely flowed through the capillaries more efficiently resulting in greater infarct targeting (Figure 4b).

These data suggested that the addition of a targeting moiety shields the untargeted self-assembling peptide nanofiber backbone from significant lung retention shortly after intravascular delivery. With the T4C-targeted nanofiber, evidence suggested a length dependence on nanofiber flow through the lung capillaries resulting in a length dependent retention within the infarct. Therefore, targeted shorter nanofibers with a narrower distribution exhibited the greatest infarct targeting.

2.3. Targeting moiety and shape dependence on retention of self-assembling peptide nanofibers

To determine the shape dependence and targeting peptide impact on retention of the nanofibers post-MI, we used the short nanofiber design since it targets the infarct most efficiently. Additionally, we continued to examine the short T4C-targeted nanofiber since it showed promising accumulation within the infarct at 1 hour. As a control, we also used a T4C-scramble targeting peptide as previously developed²² (Table 1), which was successfully incorporated onto the untargeted nanofiber backbone, and exhibited a reduced length distribution after probe sonication as visualized and quantified via TEM (Figure S9a, d). For shape dependence, a T4C-targeted and T4C-scramble self-assembling globular controls were fabricated using the previously reported globular sequence: KKQLQLQLKK²¹. This particular control was selected since the flanking charges were the same as the nanofiber and the self-assembling portion matches the ABA structure of the KQL6 nanofiber. Additionally, the self-assembling core of the globular control has half of the number of repeat units as the nanofiber therefore resulting in weaker hydrophobic interacts and a globular structure. The globular controls were imaged via TEM, and their sizes were quantified (Figure S9b, c, e) (T4C-targeted globular: 17 ± 6 nm and T4C-scramble globular: 16 ± 6 nm). All four peptides were subject to cytocompatibility evaluation, showing no adverse effects up to the expected blood concentration level post-delivery of 50 μM (Figure S6c-d). These data showed the modularity of the self-assembling peptide backbone. The nanofiber formation was not impeded by various targeting peptide additions, and the shape of the nanofiber can be modulated by changing the number of core repeat units. Additionally, probe sonication ubiquitously reduced nanofiber size distribution regardless of targeting peptide addition. It is important to note that not all targeting peptides can be added. In particular, both the T4C-targeting and T4C-scramble peptides were charged matched to the KQL6 backbone since the targeting peptide charge must not interfere with the molecular frustration generated by the lysine residues flanking either side of the hydrophobic core. When using a backbone with a different residue, this must be considered.

Nanofibers and globular controls were delivered, and the hearts (Figure 5) and satellite organs (Figure 6) were harvested, fluorescently scanned, and quantified. The globular control peptides were significantly different than the nanofibers and had the highest retention 1 hour after injection regardless of the targeting peptide addition (Figure 5a, b). The targeted nanofibers were also not significantly different than untargeted nanofibers at 1 hour (Figure 5b). This suggested that the initial self-assembling peptide accumulation within the infarct at 1 hour was not impacted by the targeting peptide sequence. At 1-day, there was a decrease in globular control infarct signal to levels more comparable to the nanofiber. This suggested that the globular control particles diffused into the infarct more easily at 1 hour. However, the true binding to available type IV collagen was saturated and the excess signal generated by increased diffusion seen at 1 hour was cleared by 1 day (Figure 5b-c). There were also significant differences between the T4C-targeted and T4C-scramble nanofibers and globular controls at 1 day (Figure 5c) suggesting that the T4C-targeting peptide specifically bound and increased retention.

Figure 5. Impact of shape and targeting peptide on self-assembling peptide left ventricle retention.

a) The left ventricle retention targeted self-assembling peptide variations. The quantified fluorescent retention of the self-assembling peptides within the left ventricle at b) 1 hour and c) 1 day. For all statistics, one-way ANOVA with Tukey’s post hoc. ns p >0.05, *p < 0.05, ** p < 0.01, *** p < 0.001, **** p < 0.0001.

Figure 6. Impact of nanofiber size on self-assembling peptide satellite accumulation 1 hour and 1 day after injection.

a) The biodistribution of the type IV collagen (T4C) targeting globular control and nanofibers vs. the T4C-scramble variants. b) The quantified fluorescent retention of the nanofibers within the satellite organs. For statistics, two-way ANOVA with Tukey’s post hoc. ns p >0.05, * p < 0.05, ** p < 0.01, *** p < 0.001, **** p < 0.0001 with comparison only performed within each satellite organ.

Analysis of the satellite organs suggests that the globular control peptides were cleared renally (Figure 6a-b). It is likely the peptides were disassembling within the bloodstream after intravascular delivery since previous work has shown that nanoparticles below 10 nm are renally cleared³². The globular peptides had average diameters over 15 nm suggesting that the assembled form of the globular controls could not be renally cleared. Disassembly could have occurred due to the weak random coil self-assembly mechanism – with a lack of β-sheet formation²¹ – previously shown for the KKQLQLQLKK sequence. With the weaker hydrophobic interactions and hydrogen bonding, this generated a random coil conformation within the globular controls and was likely insufficient to maintain the globular conformation within the shear flow of the vasculature leading to disassembly into smaller peptide components satisfying the sub-10 nm criteria for renal clearance. This also provided evidence that the nanofibers maintained their shape within the vasculature likely due to sufficient hydrophobic interactions and β-sheet formation of the six-repeat unit amino acid core.

Interestingly, the T4C-targeted nanofiber had significantly greater retention within the lungs at 1 hour compared to all the other treatments (Figure 6b). This showed that the mechanism of lung retention for the T4C-targeted nanofiber was due to increased non-specific interactions of the T4C-targeting peptide with the lung vasculature, which was fundamentally different than the untargeted nanofiber lung retention. The difference in the lung retention between the T4C-targeted nanofiber and T4C-targeted globular control suggested that the nanofiber shape increased the number of interactions with the lung vasculature via high avidity. When taken out to 1 day, there were insignificant differences in lung retention. Additionally, the T4C-scramble nanofibers were not retained within the lungs at 1 hour, which suggested that the nanofiber shape was not the sole cause of lung retention. Thus, the combination of the nonspecific binding of T4C-targeting peptides and the high avidity from a nanofiber shape generated the higher lung retention at the 1-hour timepoint.

Seeing the unique clearance behavior of the various peptides, we investigated the relative amount of remaining nanofiber and globular control within the blood as serum half-life of the peptides (Figure S10). To do so, we performed a tail vein injection of the fluorophore conjugated peptides in the same acute MI model and collected the blood at various timepoints once before tail vein injection and periodically within 60 minutes. From the blood samples, we isolated the serum and measured the fluorescence of the serum. The fluorescence of the serum is a surrogate measure for the amount of free nanofiber or globular control still circulating within the bloodstream. We plotted the fluorescence values overtime relative to the fluorescence before peptide injection and calculated the serum half-life for each peptide. The globular controls had a serum half-lie of 18.7 ± 5.8 mins, the untargeted nanofibers at 21.0 ± 4.3 mins and the T4C-targeted nanofiber at 19.5 ± 4.0 mins, all of which were not statistically different from each other (Fig S10b). This suggested that the nanofibers retained within the infarct at 1 day were due to specific targeting that occurred within an hour shortly after delivery and not from prolonged circulation. This also suggests that for the globular control, the binding sites were likely saturated within 1 hour of injection and clearance of diffusion driven accumulation was only seen at 1 day. Additionally, this data suggested that the nanofiber shape and surface ligand did not significantly impact the self-assembling peptide serum clearance.

2.4. Nanofibers target infarct vasculature while globular controls diffuse into the infarct

To investigate the morphology of nanofiber and globular control accumulation within the infarct, heart sections were stained for endothelial cells (Figure 7, S11). Overall, the short T4C-targeted nanofiber showed colocalization with endothelial cells (EC) within the infarct (Figure 7b, e). This suggested that the nanofibers were either too large or lacked sufficient conformational mobility to diffuse past the leaky vasculature generated by MI. The T4C-targeted globular control peptides showed some colocalization with the endothelial cells (Figure 7a, d,) but predominantly diffused into the infarct area. This suggested that the T4C-targeting peptide led to binding and retention in both the exposed leaky vasculature and the infarcted myocardium. This also showed that the globular controls were sufficiently small and conformationally mobile enough to diffuse past the leaky vasculature, which may be aided by the hypothesized disassembly within the blood. The T4C-scramble nanofibers and T4C-scramble globular controls (Figure S10) showed similar behavior to the T4C-targeted versions at the 1-hour timepoint (Figure 7), which further showed that initial accumulation was not impacted by targeting peptide decoration. This mechanism of infarct leaky vasculature binding for nanofibers and diffusion into the infarct was retained at 1 day (Figure S12). The overall signal intensity was lower for the globular controls while the nanofibers were comparable.

Figure 7. Immunofluorescent staining of the targeted nanofibers and globular controls within the left ventricle.

Whole heart sections with ROIs in white boxes and inset images of (a, d) type IV collagen (T4C) targeted globular control treated, (b, e) short T4C-targeted nanofiber treated and (c, f) saline treated hearts 1-hour post-MI. Scale bars for (d-f) is 100 μm. White arrows show points of colocalization with endothelial cells and white circles indicate diffusion into the infarct.

To investigate the colocalization of the T4C-targeted nanofiber to type IV collagen, 1-day post-MI heart sections were co-stained for endothelial cells and type IV collagen. Confocal microscopy performed on 1-day post heart sections (Figure 8) showed the T4C-targeted nanofiber spreading across the vasculature while partially colocalized with both endothelial cells and type IV collagen, providing evidence for specific binding. Nanofibers were found within the endothelial cell layer suggesting targeting to the leaky post-MI vasculature.

Figure 8. Confocal imaging of targeted nanofibers within the gaps of leaky vasculature within the left ventricle infarct at 1–day post-MI, colocalized with type IV collagen.

a-b) Merged confocal z-stack projections. White dotted circles indicate example areas of colocalization. Split fluorescence channels of the inset designated with the white square from panel (a) with c) nuclei in blue, d) isolectin for endothelial cells in green, e) type IV collagen in cyan and the f) short T4C-targeted nanofibers in magenta at 1-day post-MI.

These immunofluorescent staining reveal the relationship between nanomaterial shape and targeting location with the tissue. Nanofibers target leaky vasculature while globular morphologies diffuse into the infarct, which supports the use of anisotropic materials for atherosclerotic plaque targeting¹⁸, but is unique to previous work showing the use of anisotropic materials for tumor penetration^10,33. This highlights the differences in the leaky vasculature generated by MI compared to other disease models.

In addition to the microstructural differences, further work investigating this nanofiber is required since here, only a positively charged self-assembling peptide backbone. Peptide based therapeutics have shown immunogenicity often related to the charge and number of cellular interactions generated due to this charge^34,35. While negatively charged peptides have also been previously described, neutral self-assembling peptide nanofibers have been developed to address this issue of immunogenicity^30,31. Additionally, our current positively charged platform restricts the range of functional targeting ligands that can be incorporated without affecting assembly since the T4C-targeting peptide is charge matched to the backbone. Moreover, while we have shown the potential for the self-assembling peptide to be a targeted platform for post-MI treatment, a therapeutic moiety was not evaluated. Amongst the many therapeutics for MI treatment, incorporation of therapeutic peptides onto the self-assembling backbone, in the same fashion as the T4C-targeting peptide was, would be simplest. Incorporation of these peptides must obey the charge alignment with the self-assembling portion to not impeded self-assembly. While therapeutic peptides involve the simplest form of therapeutic incorporation, a myriad of chemical conjugation schemes exist, and thus non-peptide drugs may also be possible. Additionally, fluorophore conjugation also represents a model for drug conjugation. Finally, the post-MI leaky vasculature is also highly dynamic; endothelial dysfunction creates temporally varying gaps that are significantly reduced 7 day post-MI²⁰ and could alter nanofiber retention and distribution, meaning that delivery timing will likely substantially influence outcomes.

3. Conclusion

In this study we present a modular, self-assembling peptide platform that produces nanofibers of controllable length through probe sonication and ambient-temperature aging. By altering the peptide backbone, we can dictate nanomaterial shape, and the addition of defined targeting ligands enables tunable in vivo localization and retention after MI. Using compositionally similar nanomaterials, the anisotropic nanofibers preferentially bind to the gaps between endothelial cells in the post-MI vasculature through type IV collagen targeting. This behavior mirrors the unique localization to permeable vasculature of the nanofibrillar component of iECM⁹, further suggesting that the anisotropic nature of nanofibers combined with basal lamina ECM targeting leads to this unique regional targeting. In contrast, the globular nanofiber controls diffuse freely into the infarct tissue. This shape-dependent targeting provides two distinct strategies for delivering peptide therapeutic payloads to separate features of the same injury site: (1) vascular-localized delivery via nanofibers, and (2) myocardial delivery via globular nanomaterials. Our findings demonstrate that rationally designed peptide-based nanomaterials can be rapidly tuned for selective vascular versus infarct targeting, opening new avenues for shape-guided intravascular therapies after MI.

4. Methods

4.1. Self-assembling peptide synthesis and sequences

Peptides were purchased from GenScript. All peptides contained N-terminal acetylation, C-terminal amidation, TFA removal, endotoxin control and >90% purity. Peptide sequences are listed from N to C terminus in Table 1. Purification and quality control performed by GenScript.

4.2. Self-assembling nanofiber and spherical control fabrication

For all studies including characterization, in vitro and in vivo, peptides were dissolved with 0.9% saline (154 mM) sterile solution at a peptide concentration of 500 μM. The pH was adjusted to 7.4 using 0.1 M NaOH. Peptide solutions were diluted as needed using the same 0.9% saline (154 mM) sterile solution and pH was adjusted to 7.4 using 0.1 M NaOH as needed.

For sonication and filtration studies, the peptide solutions were prepared as above and then immediately probe sonicated on a 500W 20 kHz ultrasonic probe sonicator (Fisherbrand) with a 1/16” tip. The optimized amplitude was 30% for 10 minutes of active sonication in cycles of 15 seconds on and 45 seconds off while keeping the tube fully submerged in wet ice to reduce sample heating and excess evaporation. Sonicated nanofibers were then filtered using a 4mm 0.22 μm PES syringe filter (Restek). Nanofibers were subject to characterization, fluorophore conjugation for same day in vivo use for 1 hour or aged up to 21 days after for analysis. Non-sonicated nanofibers and globular controls were used immediately for analysis or conjugation with a fluorophore after pH adjustment.

4.3. Transmission Electron Microscopy

For TEM, 200 mesh, Formvar carbon copper TEM grids (EMS) were glow discharged, inverted onto droplets containing nanofibers for 30 seconds. Globular control samples were incubated for 2 minutes. Grids were inverted onto water droplets 3 times for 30 seconds and stained with 2% uranyl acetate for 1 minute and finished with a 4-second blot. Samples were air-dried before imaging on the Jeol 1400 plus operated at 80 KeV and equipped with a bottom-mounted Gatan One View camera. To measure the nanofiber distribution sufficiently, 3 or more images were taken per condition. Each image was divided into a quadrant and a minimum of 50 nanofibers per quadrant were measured manually using Fiji/ImageJ using the “Freehand Lines” tool. Multiple images were analyzed with a minimum of 1000 fibers measured per condition. For the globular controls, a minimum of 200 structures were measured using the “Straight Line” tool.

4.4. In vitro cytocompatibility

Cytocompatibility of the nanofiber and spherical controls was measured using L929 murine fibroblasts. Fibroblasts were seeded at 10k cells per well in a 96 well plate and allowed to adhere and proliferate for 24 hours. After 24 hours, a baseline viability was quantified via Alamar Blue. Peptide nanofibers and spherical controls were prepared as in Section 4.2. These peptides along with a 0.1% zinc diethyldithiocarbamate (ZDEC) control in 1x PBS and saline control were doped into each well. The peptides were doped into each well after serial dilution with saline and pH 7.4 adjustment with 1M NaOH and allowed to incubate for 24 hours after which an endpoint reading was quantified by Alamar Blue. The Alamar Blue was allowed to incubate for 3 hours prior to absorbance measurement at 570 nm on a plate reader (Biotek Synergy HTX).

4.5. Osmolarity, colloidal oncotic pressure and hemolysis percentage

The osmolarity, colloidal osmotic pressure and hemolysis were measured using methods adapted from previously³⁶. Osmolarity was measured using a freezing point osmometer (Precision Systems, Model 6002; Natick, MA). The colloidal osmotic pressure was determined using a membrane colloid osmometer (model 4420, Wescor; Logan, UT).

To measure the % of hemolysis, we used fresh whole blood purchased from San Diego Blood Bank. Blood from 5 different donors was used. The red blood cells (RBCs) were isolated by centrifuging 40 mL of whole human blood at 800 rpm for 10 min, followed by gentle aspiration of the buffy coat. After gentle mixing of the blood, 1 mL of blood was transferred to a sterile tube and mixed with 100 μL of peptide solution to obtain a 1:10 dilution (since we only have a small volume of peptide in some cases, we mixed 0.5 mL of blood with 50 uL of peptide solution). The blood mixtures were gently mixed and incubated at 37 °C for 1 hour before measuring the % hemolysis. For each sample and replicate, the total hemoglobin and the plasma-free Hb were measured before and after ultracentrifugation using a capillary microhematocrit (Clay Adams Micro-Hematocrit; Farmingdale, NJ). Total Hb and plasma-free Hb content were determined spectrophotometrically (B-Hemoglobin, Hemocue, Stockholm, Sweden). To calculate % hemolysis, the plasma-free Hb was divided by the total Hb and reported as a percentage. Physiological saline with 0.5% human serum albumin (AlbuRx 25, CSL Behring, King of Prussia, PA) was used as a positive control, and 1% Triton X-100 (Sigma Aldrich) in distilled water was used as a negative control.

4.6. Rat model of MI and I/R Injury

All procedures in this study were approved by the Committee on Animal Research at the University of California San Diego and in accordance with the guidelines of the Association for the Assessment and Accreditation of Laboratory Animal Care (A3033-01). Adult female Sprague-Dawley rats (225-250g) were used to induce MI as previously described³⁷. The coronary artery was accessed via thoracotomy and was ligated for 35 minutes. After the ischemic period, the suture was released to restore blood flow.

4.7. Biodistribution and serum half-life of the nanofiber and spherical controls.

For biodistribution and LV localization of the nanofibers and globular controls, the peptides were fluorophore conjugated after dissolution and pH adjustment as mentioned in Section 4.2. All conjugations were conducted at a peptide concentration of 500 μM and all tail vein injections were performed using the 500 μM conjugated peptide solution. A 1:200 volume ratio of 10 mg/mL IVISense 750 NHS Fluorescent Self-Quenching Dye (Revvity) was added to the dissolved peptide solution and allowed to conjugate for 1 hour at 20°C prior to tail vein injection. After reperfusion, the nanofiber and globular controls were delivered via tail vein injection at 500 μM and 1 mL volume using a 27G needle (BD) and 3 mL syringe (BD). Control animals received 1 mL of sterile saline. At 1 hour and 1-day post-injection, rats were euthanized and organs collected for ex vivo imaging on an Odyssey CLx (LICORbio) in the 800-fluorescence emission channel. Heart and satellite organs (lungs, kidneys, spleen, liver) were extracted, with the heart sliced into 4-6 short-axis slices. Average signal intensity of the left and right ventricles and organs were measured using the LICORbio Image Studio software and figure images were exported.

For serum half-life of the peptide nanofibers and globular control, acute MI surgery in Section 4.5 was first performed followed by tail vein injection of the fluorophore conjugated peptide nanofibers or globular control. Blood was collected prior to tail vein injection of the peptides, and at 1, 15, 30, 45 and 60 minutes after tail vein injection. Extracted blood was allowed to coagulate at room temperature for 20 minutes before centrifugation for 10 minutes at 20,000g. The serum was scanned on an Odyssey CLx (LICORbio) using the 800-fluorescence emission channel. The percentage of fluorescence signal was calculated with the 1-minute serum value as the maximum. The serum prior to tail vein injection yielded <0.01% signal relative to 1-minute post, and the serum 1 minute after injection was considered 100%. Serum was collected in biological triplicate, scanned in technical triplicates and averaged.

4.8. Immunofluorescence staining and imaging

Heart sections harvested were frozen in Tissue-Tek O.C.T. Compound (Sakura) and sectioned into 10-micron slices mounted on glass slides. Sections were stained using DAPI, GSL I isolectin B4 fluorescein (Vector Labs - FL-1201-.5) and type IV collagen (Sigma SAB4200500) and imaged on a Slideview VS200 (Olympus). Images were analyzed and exported using QuPath. These slides were also subject to confocal imaging at 40x on an Andor Dragonfly 200 (Oxford Instruments).

4.9. Statistical Analysis

Statistical analysis was performed on GraphPad Prism10. For all plots, mean ± standard deviation is plotted. For all nanofiber measurements, more than 400 nanofibers were measured total per sample across n = 5-8 TEM images. For all in vivo analysis, sample size was n = 3-4 biological replicates. For analysis of multiple groups, a one-way ANOVA with post hoc Tukey’s test was performed. For single comparisons, a two-tailed T-test was performed. Significance was accepted at p < 0.05. For half-life calculation, the data was fitted to a one phase exponential decay: y_o set to 100 and the equilibrium set to 0.