“Missing Tooth” Multidomain Peptide Nanofibers for Delivery of Small

Article pubs.acs.org/Biomac

“Missing Tooth” Multidomain Peptide Nanofibers for Delivery of Small Molecule Drugs I-Che Li, Amanda N. Moore, and Jeffrey D. Hartgerink* Departments of Chemistry and Bioengineering, Rice University, 6100 Main Street, Houston, Texas 77005, United States S Supporting Information *

ABSTRACT: The clinical administration of many small molecule hydrophobic drugs is challenged by the insolubility of these drugs under physiological conditions. Because of this, the development of biocompatible scaffolds capable of effectively delivering hydrophobic drug molecules is of particular interest. Multidomain peptides (MDPs) provide biocompatible hydrogel scaffolds that are injectable and spaceconforming, allowing for in situ delivery of a variety of drugs. Here we demonstrate that through manipulation of peptide primary sequence, a molecular cavity can be incorporated into the hydrophobic core of these peptide nanofibers allowing for encapsulation and delivery of small molecule drugs with poor water solubility. Using SN-38, daunorubicin, diflunisal, etodolac, levofloxacin, and norfloxacin, we demonstrate drug encapsulation and release from multidomain peptide fibers. Steady-state fluorescence and drug release studies show that hydrogels loaded with SN-38, diflunisal, and etodolac exhibit prolonged drug release profiles due to intrafibrillar drug encapsulation. This study establishes multidomain peptides as promising carriers for localized in situ delivery of small molecule drugs with poor water solubility.

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INTRODUCTION The field of drug delivery strives to develop biocompatible materials as carriers of clinical therapeutics.1,2 For clinical applications of many small molecule drugs, water insolubility poses a significant challenge in the administration of these drugs.3,4 Lipid-based micelles, microparticles and emulsionbased techniques have previously been used to encapsulate hydrophobic drugs for delivery to the body.5 While these types of strategies circumvent issues with drug insolubility, they fail to address the need for localization of drug activity. Because these drugs are introduced systemically, the entire body is treated with the therapeutic agent rather than simply the diseased area, thereby resulting in unnecessary and undesirable side effects.6−8 Therefore, localized drug delivery has become important for drug delivery. Minimizing side effects through localized and site-specific targeting of drugs is a key objective.6 To accomplish this, development of vehicles capable of localizing delivery and controlling drug release kinetics is paramount.9,10 A variety of polymeric hydrogels that mimic the native extracellular matrix of the body have been used in tissue engineering strategies.11−15 Specific physical properties of hydrogels, including injectability and shear recovery also establish these materials as potential candidates for successful loading and localized delivery of drugs.16,17 Due to the easy manipulation of chemical and physical properties, by changing the amino acid building blocks, peptide-based supramolecular biomaterials specifically have been widely developed to solvate various © XXXX American Chemical Society

agents, such as carbon nanotubes, growth factors, and small molecule drugs.18−22 Recently, several injectable supramolecular biomaterials have been applied to achieve localized drug delivery. The “RADA16” peptide hydrogel has been reported to nonspecifically encapsulate a variety of bioactive proteins without changing their functionality.22 Through passive diffusion, functional proteins are released from the gel.23 For noncovalent drug encapsulation, the “MAX” series of peptides which form hydrogels with high storage moduli have been used.24,25 For example, “MAX8” has been used to encapsulate curcumin, a hydrophobic chemotherapeutic drug, primarily through hydrophobic interactions.26 Peptide amphiphile (PA), a widely applied supramolecular biomaterial, has been reported to carry various molecules by modulating the building blocks.27 Through host−guest interaction, the PA nanofibers can encapsulate small molecules within the hydrophobic core.28 Covalent attachment has also been used to deliver anticancer drug molecules. In this strategy, drug molecules are conjugated to the filamentous supramolecular peptides using a linker located at the N-termini of peptides.29 When the peptide selfassembles in a buffer solution, the peptide-drug conjugates form a hydrogel, which acts as a highly efficient drug carrier. Received: March 1, 2016 Revised: May 20, 2016

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DOI: 10.1021/acs.biomac.6b00309 Biomacromolecules XXXX, XXX, XXX−XXX

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Figure 1. Schematic of missing tooth multidomain peptides. (A) Missing tooth MDP sequence K2(SL)2(SA)2(SL)2K2. (B) MDP with specific modification to leave a gap in the packing of hydrophobic amino acids in the core of the self-assembled fiber. This gap can be filled with small molecules with the appropriate size and chemistry. (C) Drug molecules stay in the exterior of hydrogel nanofibers allowing for rapid diffusion of drugs from the gel. (D) Drug molecules exhibit prolonged drug release due to slow release from nanofibers and the following diffusion from the gel.

nanofibers with a hydrophobic core and hydrophilic surface. However, the accumulation of positive charges along the fiber (due to lysine side chains) results in short peptide nanofibers. By adding multivalent ions of opposite charge, the electrostatic repulsion of termini is shielded allowing for long-range fiber growth. Entangling and noncovalent cross-linking of fibers ultimately gives rise to a three-dimensional hydrogel scaffold.31 Our group has previously used MDP hydrogels for controllable release and delivery of growth factors, cytokines, and multivalent ionic molecules. For instance, the growth factors VEGF, FGF2, and TGFβ1 were incorporated into a K(SL)3RG(SL)3KGRGDS hydrogel through heparin binding.32 Another study by our lab established a method for the bimodal release of growth factors and cytokines from the MDP hydrogel.20 In an additional study, we demonstrated increased macrophage infiltration and polarization due to biphasic cytokine release.33 A final study evaluated MDP hydrogels as

In order to deliver small molecule drugs with poor water solubility, here we developed a specifically engineered material to hold drugs as cargo. In this study, we designed injectable multidomain peptide (MDP) hydrogels that have a hydrophobic binding pocket to store hydrophobic molecules. In contrast to liquid-based carriers, the MDP hydrogel will remain localized to the site of injection allowing site-specific administration of the drug. MDPs incorporate a base peptide design of alternating hydrophobic and hydrophilic amino acids and charged residues at both termini.30 In aqueous solutions, the hydrophobic side chains will aggregate to form a “sandwich” peptide dimer. A fiber consisting of multiple peptide dimers forms to further sequester hydrophobic residues from contact with the surrounding aqueous solution, and this nanofiber is stabilized by hydrogen bonds that form between the peptide backbones of adjacent dimers. Therefore, self-assembly results in β-sheet B

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solutions for TEM, lyophilized peptides were dissolved to 2 wt % in Milli-Q water and subsequently diluted 50:50 with 1× Hank’s buffered saline solution (HBSS; Life Technologies). Serial dilutions were then performed with Milli-Q water to reach a peptide concentration of 0.01 wt %. The diluted sample solutions were pipetted onto a Quantifoil R1.2/1.3 holey carbon mesh copper grid and allowed to sit for 1 min. Excess solution was wicked away with filter paper and the grid was negatively stained with 2 wt %, pH 7 phosphotungstic acid (PTA) for 5 min. Excess stain was wicked away from the grid before drying overnight. A 100 kV JEOL 2010 transmission electron microscope (JEOL USA Inc., Peabody, MA) was performed to image the dried grids. For atomic-force microscopy (AFM) measurement, the sample solutions were prepared and diluted in the same manner as the TEM samples. Each diluted sample was spin-coated onto a freshly cut mica disc. Eight microliters of sample solution was dropped onto the mica surface, allowed to dry for 8 s, and then rinsed with Milli-Q water for 10 s to remove salt crystals. The sample was then spun dry for 10 min using a Headway Research photoresist spinner. The dried sample was imaged in tapping mode using a Digital Instruments Nanoscope IIIa. Rheological Analysis. 1× Dulbecco’s phosphate-buffered saline (DPBS) was prepared by adding powdered medium to 9.55 g/L in Milli-Q water, and adjusting the pH to 7. MDPs were dissolved in 1× DPBS to form 1 wt % hydrogels. To evaluate the rheological properties of K2(SL)3SA(SL)2K2 and K2(SL)2(SA)2(SL)2K2 hydrogels, the storage modulus (G′) and loss modulus (G″) were monitored during strain sweep and shear recovery experiments performed using a TA Instruments AR-G2 rheometer (TA Instruments, New Castle, DE). One hundred fifty microliters of prepared hydrogel was deposited onto the rheometer stage, and a 12 mm stainless steel parallel plate was used with a 1000 μm gap height. In the strain sweep analysis, G′ and G″ were monitored under an applied strain of 0.01% to 200% at a frequency of 1 rad/s. In the frequency sweep analysis, G′ and G″ were monitored under 1% strain at frequency of 0.1 rad/s to 100 rad/s. Shear recovery experiments were performed by subjecting the gel to 1% strain for 20 min, increasing the strain to 200% for 1 min, and then reducing the strain back to 1% for 20 min. Drug Encapsulation. SN-38, daunorubicin, and etodolac were purchased from VWR International (Radnor, PA); diflunisal, levofloxacin and norfloxacin were purchased from Sigma-Aldrich (St. Louis, MO). Daunorubicin, etodolac, diflunisal, levofloxacin, and norfloxacin were dissolved in DMSO at a concentration of 0.2 wt % and then added to 2 wt % MDP solutions in Milli-Q water. SN-38 was dissolved in DMSO at the concentration of 0.005 wt % and then added to MDP solution because of its low solubility. After the addition of drugs, concentrated (38.2 g/L) DPBS buffer was added to prepare 1 wt % MDP hydrogels with 10% DMSO in 1× DPBS buffer. The concentration of SN-38 in the hydrogel is 0.0005 wt %. The concentration of other drugs in the hydrogels is 0.02 wt %. After 24 h, the fluorescence of the drug-carrying hydrogels was tested using a TECAN Infinite 200 plate reader (Tecan US, Inc., Morrisville, NC). Drug Loading and Release. In order to test the potential use of the target MDP hydrogels for drug delivery, the drug-carrying MDP hydrogels were prepared as above and dialyzed. The hydrogels were loaded into a sealed dialysis bag and then dialyzed against DPBS, with gentle shaking by a test tube rocker during dialysis. As a control, drugs alone were dissolved to 0.02 wt % in 1× DPBS buffer, and also dialyzed against DPBS. To monitor the amount of drug released into DPBS buffer outside of the dialysis bag, 100 μL of the of the outer liquid was removed, and the fluorescence was measured using plate reader at 1, 2, 4, 8, 12, 24, 48, 72, 96, 120, 144, and 192 h time points. The fluorescence value at each time point was calibrated using a standard curve to give drug concentration, and these concentrations were used to calculate the cumulative release amount of the drug.

a long-term delivery agent for a variety of clinically relevant multivalent molecules used to cross-link the scaffold.19 Expanding MDPs to allow delivery of hydrophobic small molecules from the MDP hydrogel is complicated by the aqueous nature of the hydrogel; therefore, a novel strategy to accomplish drug delivery by creating a binding site in the hydrophobic core of MDP fibers is required. To effectively encapsulate the drug in the nanofiber interior, a stable binding structure must be considered when designing the MDP sequence. Here we report a new material designed to controllably deliver small molecules with low water solubility. In this design, we use the previously designed K2(SL)6K2 MDP as a template and substitute the middle one or two leucine residues with alanine residues to make K2(SL)3SA(SL)2K2 and K2(SL)2(SA)2(SL)2K2, respectively. For convenience, we define peptide sequence K2(SL)6K2 as “SL”, K2(SL)3SA(SL)2K2 as “SL1” and K2(SL)2(SA)2(SL)2K2 as “SL2”. By substituting the core hydrophobic residues for those with smaller amino acid side chains (Figure 1A), MDPs will self-assemble to form defective sandwich dimers (Figure 1B). When considering an entire nanofiber consisting of several defective dimers, the defects align to generate a cavity along the length of the nanofiber. We hypothesize that the cavity created in the hydrophobic core of the nanofibers will stabilize entrapment of hydrophobic drugs through the hydrophobic effect. This design, we call a “missing tooth” design because of the shape. These peptides were utilized to construct hollow fibers for the encapsulation of three groups of drugs, anticancer drugs (daunorubicin, SN-38), nonsteroidal anti-inflammatory drugs (diflunisal, etodolac) and antibiotics (levofloxacin, norfloxacin). The three groups of drugs have specific goals of treatment, such as cancer cell removal, wound healing, and infection prevention, that we hope to extend the potential of delivery with missing tooth design. In addition, the fact that all the drugs we chose are fluorescent will benefit the study of their intrafibrillar encapsulation. By using the fluorescence of these molecules as a probe, the gels exhibit intrafibrillar encapsulation and controlled release of hydrophobic drugs (Figure 1C,D). With the “missing tooth” design, the MDP hydrogel can be used as an effective carrier of small molecules with low water solubility.

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EXPERIMENTAL METHODS

All chemicals not otherwise specified were purchased from SigmaAldrich (Sigma-Aldrich, St. Louis, MO). Peptide Synthesis. K2(SL)3SA(SL)2K2 and K2(SL)2(SA)2(SL)2K2 were synthesized following a previously described solid-phase peptide synthesis protocol, using Rink amide resin with 0.37 mM loading and N-terminal acetylation.31 All resin and coupling reagents were purchased from EMD Chemicals (Philadelphia, PA). After cleavage from the resin, peptides were purified by dialysis for 7 days using 1000 Da MWCO dialysis tubing (Spectra/Por, Spectrum Laboratories Inc., Rancho Dominguez, CA) against Milli-Q water. The dialyzed peptide solutions were frozen and lyophilized. All peptides were characterized by matrix assisted laser desorption/ionization time-of-flight mass spectrometry, Autoflex MALDI-ToF MS (Bruker Instruments, Billerica, MA), to verify a successful synthesis. Secondary Structure. To investigate the secondary structures of the synthesized peptides, attenuated total reflectance Fourier transform infrared spectroscopy (ATR FT-IR) and circular dichroism (CD) were used. In ATR FT-IR assessment, a 0.1 wt % peptide solution was allowed to dry on the diamond of a “Golden Gate”. The CD spectra were recorded from 180 to 250 nm using a Jasco-810 spectropolarimeter. Nanostructure. Negative-stain TEM was performed on both MDPs to confirm nanofiber formation. To prepare gelled sample

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RESULTS AND DISCUSSION Secondary Structure. K 2 (SL) 3 SA(SL) 2 K 2 and K2(SL)2(SA)2(SL)2K2 were synthesized by standard Fmoc based solid phase peptide synthesis, details for which are C

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Figure 2. Secondary structural characterization of missing tooth structures. FT-IR spectrum of (A) SL1: K2(SL)3SA(SL)2K2 and (B) SL2: K2(SL)2(SA)2(SL)2K2 showing characteristic peaks for β-sheet: 1630 cm−1 and antiparallel: 1695 cm−1. CD spectrum of (C) SL1 and (D) SL2 showing characteristic maximums at 195 nm and minimums at 216 nm for β-sheets.

included in the Supporting Information. Their respective masses were verified by MALDI-TOF MS (spectra shown in Supporting Information). Both ATR FT-IR and CD were performed to assess the secondary structures of the designed MDPs. Since K2(SL)3SA(SL)2K2 and K2(SL)2(SA)2(SL)2K2 are designed based on K2(SL)6K2, they were expected to show similar secondary structure in both techniques.31 As seen in the IR spectra shown in Figures 2A and 2B, both peptides exhibited an amide I peak near 1630 cm−1, which indicates a β-sheet secondary structure.34 Also, the shoulder peaks near 1695 cm−1 suggest that both of the peptides form β-sheets with an antiparallel orientation, which is consistent with previously reported K2(SL)6K2.30,31,35 In the CD results, shown in Figures 2C and 2D, both peptides have a maximum at 195 nm and a minimum at 216 nm, which can also be correlated to β-sheet structure.31 Thus, the results of both IR and CD confirmed that K2(SL)3SA(SL)2K2, K2(SL)2(SA)2(SL)2K2, and K2(SL)6K2 all adopt an antiparallel β-sheet secondary structure. This demonstrates that the substitution of alanine for leucine in these positions does not significantly alter peptide secondary structure. Nanostructure. To confirm the self-assembly and nanofiber formation of the target MDPs, negative-stain TEM was performed. In Figure 3A,B, the images show that both of the missing tooth MDPs form nanofibers with similar morphology to those of K2(SL)6K2.31 These images of K2(SL)3SA(SL)2K2 and K2(SL)2(SA)2(SL)2K2 nanofibers verify that the substitution of middle hydrophobic amino acids is not prohibitive of nanofiber self-assembly. For further confirmation by AFM, as seen in Figure 3C,D, both of the target peptides form nanofibers similar to previously developed MDPs, with a height of 2 nm (see Supporting Information for details).30 In Figure 3C, K2(SL)3SA(SL)2K2 reveals a long nanofiber length, which suggests structural stability of the nanofiber despite the alanine substitution. Compared to 3C, 3D shows that the

Figure 3. Nanostructural characterization of missing tooth structures. Negative stained TEM of (A) K 2 (SL) 3 SA(SL) 2 K 2 and (B) K2(SL)2(SA)2(SL)2K2. AFM of (C) K2(SL)3SA(SL)2K2 and (D) K2(SL)2(SA)2(SL)2K2. Scale bars for A and B are 50 nm, and for C and D 500 nm. Fibrillar structure is similar to the unmodified K2(SL)6K2 peptide which exhibits fiber length 120 ± 30 nm, width 6 ± 1 nm, and height 2 nm.31

double alanine substitution in K2(SL)2(SA)2(SL)2K2 forms shorter nanofibers in general. This may be due to the double alanine substitution decreasing the stability of the nanofiber. We conclude that the hydrophobic defect in the missing tooth D

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alanine does not interrupt this property, a shear recovery experiment was performed. Under high percent strain, we anticipate an inversion of G′ and G″ values which indicates hydrogel shear thinning; however, when this high percent strain is removed, G′ should recover to its initial value. After the shear thinning event, the G′ of both K2(SL)3SA(SL)2K2 and K2(SL)2(SA)2(SL)2K2 recovered to their initial values in less than 5 min, indicating both hydrogels are capable of recovering from shear thinning (Figure 4A,B). These results confirmed that these hydrogels are injectable materials and can be used in syringe localized drug delivery studies. Drug Encapsulation. Fluorescence emission peaks of small molecules, such as SN-38, are known to shift in solvents of different polarities.38−40 Because the environment within the nanofiber cavity will be more hydrophobic than the aqueous environment outside the nanofiber, we can use fluorescence emission shifts to monitor intrafibrillar encapsulation of the drugs. Here we chose small molecule drugs, SN-38, daunorubicin, diflunisal, etodolac, levofloxacin and norfloxacin to test the ability of missing tooth MDPs to encapsulate and release small molecule drugs. We monitor the fluorescence emission spectra prior to and following gel encapsulation, and compare these spectra to those of drug molecules solvated in a variety of solvents with known polarity. The fluorescence results should indicate the hydrophobicity of the environment in which the drugs are solvated, with drugs encapsulated within the nanofibers showing a more hydrophobic environment. In order to characterize the encapsulation of small molecule drugs in hydrogel nanofibers, SN-38, daunorubicin, diflunisal, etodolac, levofloxacin, and norfloxacin were dissolved in 2 wt % MDP solutions, followed by the addition of DPBS buffer to prepare 1 wt % MDP hydrogels in 1× DPBS buffer. CD and TEM of the hydrogel after drug encapsulation were characterized and showed no significant change of secondary structure and fibrillary morphology (SI Figures 4 and 5). The solutions were then tested using a fluorescent plate reader. Figures 5A and 5D reveal the steady-state fluorescence of SN-

design may destabilize the nanofiber structure, but still allows for the self-assembly and formation of nanofibers. Rheological Properties. To evaluate the viscoelastic properties of the K2(SL)3SA(SL)2K2 and K2(SL)2(SA)2(SL)2K2 hydrogels, oscillatory rheometry was used. The analysis results under low percent strain are shown in Figure 4. For the

Figure 4. Rheological characterization of peptide shear recovery. After constant shearing of 1% for 20 min, peptides were sheared at 200% for 1 min and then shearing returned to 1% for 20 min. Recovery to >95% G′ prior to perturbation is noted after 1 min: (A) K2(SL)3SA(SL)2K2 and (B) K2(SL)2(SA)2(SL)2K2.

K2(SL)3SA(SL)2K2 hydrogel, the G′ value of 142 Pa is approximately 1 order of magnitude greater than its G″ of 13 Pa, indicating hydrogel formation. With a G′ of 158 Pa and a G″ of 14 Pa, the double-modified K2(SL)2(SA)2(SL)2K2 also forms a hydrogel. Compared with the G′ of K2(SL)6K2, which was reported as 191 Pa, the storage moduli of K2(SL)3SA(SL)2K2 and K2(SL)2(SA)2(SL)2K2 are roughly equivalent.31 The slight reduction in storage moduli is consistent with the hypothesized defect structure in the hydrophobic core of nanofibers. For clinical application of a biomaterial, ease and invasiveness of delivery prove important considerations. Previously developed MDPs have demonstrated the property of shear recovery, allowing this class of biomaterial to be injected at a site of interest.19,36,37 To verify that substitution of leucine for

Figure 5. Steady-state fluorescence emission spectra for (A) SN-38, (B) diflunisal, (C) etodolac in K2(SL)6K2 hydrogel, K2(SL)3SA(SL)2K2 hydrogel, and K2(SL)2(SA)2(SL)2K2 hydrogel. The fluorescence of (D) SN-38, (E) diflunisal, and (F) etodolac in solvents of different polarities are also given. E

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Figure 6. Steady-state fluorescence emission spectra for (A) daunorubicin, (B) levofloxacin, (C) norfloxacin in K2(SL)6K2 hydrogel, K2(SL)3SA(SL)2K2 hydrogel, and K2(SL)2(SA)2(SL)2K2 hydrogel. The fluorescence of (D) daunirubicin, (E) levofloxacin, and (F) norfloxacin in solvents of different polarities are also shown.

Figure 7. Release profiles for (A) SN-38, (B) diflunisal, (C) etodolac, (D) daunorubicin, (E) levofloxacin, and (F) norfloxacin, showing the release from K2(SL)6K2 hydrogel (blue), K2(SL)3SA(SL)2K2 hydrogel (red), K2(SL)2(SA)2(SL)2K2 hydrogel (green), and DPBS as a control (black). n = 4 for (A), (B), and (C).

38 in the three hydrogels, as well as in solvents of different polarities. In isopropanol, which is the least polar solvent tested, SN-38 has a strong emission peak in the green light region around 430 nm indicating the low polarity of the environment. When SN-38 is dissolved in methanol, water, or DPBS (in order of increasing polarity), the 430 nm emission peak decreases with increasing solvent polarity. In Figure 5A, K2(SL)3SA(SL)2K2 hydrogel (blue dashed line) has the highest emission peak of all gels at 430 nm. The lower fluorescence intensity of SN-38 in hydrogel than in solvent indicates that part of the SN-38 is encapsulated inside the hydrophobic core of SL1 nanofibers and part of it still stayed in high polarity environment. By contrast, the weak emission at 430 nm of

K2(SL)6K2 and K2(SL)2(SA)2(SL)2K2 hydrogels suggested only a small amount of or no intrafibrillar encapsulation. For diflunisal in different solvents (Figure 5E), the emission peak decreased and red-shifted from 330 to 350 nm in solvents of lower polarity, like methanol or isopropanol. In the spectra of diflunisal in hydrogels (Figure 5B), all the three hydrogels, K2(SL)6K2, K2(SL)3SA(SL)2K2 and K2(SL)2(SA)2(SL)2K2, have basically the same emission maximum. Their fluorescence spectra are similar to the DPBS control in Figure 5E. In the etodolac fluorescence study, the spectra (Figure 5F) showed that the emission intensity reduces in solvents with higher polarity, such as water and DPBS. This decrease is due to fluorescence quenching by polar solvent molecules. In Figure 5C, the spectra also revealed that the K2(SL)3SA(SL)2K2 F

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ations for their steady-state fluorescence data, while that of diflunisal does not. Figure 7B shows a delayed release of diflunisal from all three hydrogels, with differing rates of release for each hydrogel. This is inconsistent with the steady-state fluorescence data, which suggests that intrafibrilliar encapsulation of diflunisal may not be occurring (Figure 5B). One possible explanation for this inconsistency is the low fluorescence intensity of diflunisal in a hydrophobic environment. Unlike SN-38 and etodolac, the low fluorescence of diflunisal in a hydrophobic environment could be overwhelmed by the hydrophilic peak and difficult to detect. Another possible reason for the attenuated release of diflunisal is the interaction with hydrogel scaffolds on the outer surface of the nanofibers.19 For example, the carboxyl and hydroxyl groups of diflunisal could interact with serine or lysine side chains via ionic crosslinking or hydrogen bonding. To understand the interaction, future studies may be able to use time-resolved fluorescence spectroscopy to measure diflunisal fluorescence lifetime change in solvents of different viscosities or hydrophobicities.41,42 Also, nuclear magnetic resonance (NMR) spectroscopy might provide further evidence of intrafibrillar encapsulation.43,44 To explain the trends of drug intrafibrillar encapsulation, we hypothesized that the efficiency of encapsulation and release of the drugs is related to hydrophobicity. Hydrophobicity of the drug molecules can be correlated to their log P value, which is the logarithm of n-octanol/water partition coefficient. A log P values, which were obtained from the ChEMBL database, are predicted values using the atomic contribution method.45,46 Here we compare the computational A log P of the drugs in Table 1. The computational value A log P have similar trends

hydrogel solution has a greater emission intensity than the other MDPs, which suggests that more etodolac molecules are more effectively encapsulated in K2(SL)3SA(SL)2K2 nanofibers. For daunorubicin, levofloxacin, and norfloxacin, Figure 6 shows the fluorescence profiles. The spectra of the drugs look similar in all hydrogels (Figures 6A−C) while the spectra vary in different solvent environments (Figures 6D−F). Also, the fluorescence spectra in gels are almost the same to the DPBS control, suggesting poor intrafibrillar encapsulation in all hydrogels. Drug Release Profile. The quantification of drug release was performed for K 2 (SL) 6 K 2 , K 2 (SL) 3 SA(SL) 2 K 2 , K2(SL)2(SA)2(SL)2K2 hydrogels with an n = 4. Drugs dissolved in DPBS without MDP hydrogel encapsulation were also used as controls to mimic simple passive release of the drug. Each drug was solvated in the MDP solutions during gelation as described in the experimental section. By measuring the fluorescence of the dialysis buffer at each time point, the percentage of drug released into the buffer was then calculated to make the profile in Figure 7. In the missing tooth design, we hypothesize that hydrophobic drugs would be encapsulated in the nanofibers’ aliphatic core because of the hydrophobic effect. As a result, the hydrogel, which has better affinity with the drugs, may delay the drug release the most. In the steady-state fluorescence study, K2(SL)3SA(SL)2K2 shows the highest drug encapsulation ability. Here, in Figure 7A, K2(SL)3SA(SL)2K2 has the slowest release rate among all hydrogels as we predicted. Around 60% of the loaded SN-38 is released from the K2(SL)3SA(SL)2K2 hydrogel over 8 days. In comparison, hydrogels K2(SL)6K2 and K2(SL)2(SA)2(SL)2K2 released more than 90% of the loaded SN-38 in 24 h. Based on similar release curves for the DPBS control, we suggest that most of the drug molecules stayed in the exterior of hydrogel nanofibers and therefore exhibit shortterm release through passive diffusion. For diflunisal, which degrades over time, the release curves were corrected with the degradation curve, and the corrected curve is shown as Figure 7B (see Supporting Information for details). Compared to the DPBS control, all tested hydrogels showed delayed drug release. Considering the relative release rates of the three gels, K2(SL)3SA(SL)2K2, K2(SL)6K2, K2(SL)2(SA)2(SL)2K2 (in order from slowest to fastest), K2(SL)3SA(SL)2K2 hydrogel delayed the release the most, suggesting the most effective intrafibrillar encapsulation. Similarly, etodolac in Figure 7C shows delayed release in gels. The order of release rate from slowest to fastest is K2(SL)3SA(SL)2K2, K2(SL)2(SA)2(SL)2K2, K2(SL)6K2. This result is consistent with the steady-state fluorescence data, which indicates that K2(SL)3SA(SL)2K2 has the most drug encapsulation among all hydrogels. By contrast, daunorubicin, levofloxacin, and norfloxacin exhibit short-term drug release regardless of hydrogel encapsulation, as shown in the Figures 7D−F. For daunorubicin, shown in Figure 7D, all gels released more than 95% of the loaded daunorubicin within 4 h. Similarly, more than 80% of levofloxacin (Figure 7E) and 70% of norfloxacin (Figure 7F) were released from the gels within 4 h. Also, for the three drugs, all gels have similar release curves as DPBS controls. These results suggest that daunorubicin, levofloxacin, and norfloxacin tend to stay in the exterior of nanofibers rather than in the interior, allowing for rapid diffusion of drugs from the gel. Here the release profiles of the five drugs, SN-38, etodolac, daunorubicin, levofloxacin and norfloxacin, match our expect-

Table 1. Computational A log P Values and the Release of the Small Molecule Drugs

a

drug name

A log P

delayed release

SN-38 diflunisal etodolac daunorubicin levofloxacin norfloxacin

2.45a 3.15a 3.72a 0.63a −1.37a −1.41a

yes yes yes no no no

A log P values were obtained from the ChEMBL database.

with the drug encapsulation tendencies. In general, SN-38, diflunisal and etodolac have larger A log P’s than daunorubicin, levofloxacin and norfloxacin. For SN-38 and etodolac, K2(SL)3SA(SL)2K2 has the best efficiency of intrafibrillar encapsulation among the three hydrogels, which explains the slow release of drugs. In contrast, daunorubicin, levofloxacin, and norfloxacin all have low intrafibrillar encapsulation according to steady-state fluorescence, results which are also consistent with their short-term releases and match our hypothesis. Diflunisal, as an exception, is likely not encapsulated in nanofibers but has slow release. It is suggested that there could be other determinants dominating in this case, such as the interactions discussed in the previous paragraph. Since the drug hydrophobicity is the dominating determinant in 5 out of the 6 drugs, the A log P value can still be a useful tool to explore potential small molecule drugs for long-term delivery.

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CONCLUSIONS In this report, we have demonstrated the ability of “missing tooth” MDP hydrogels to encapsulate and release small G

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molecule drugs with low water solubility. Specifically, the hydrophobic defect in the missing tooth design enhanced the solvation of SN-38, diflunisal, and etodolac in the aliphatic core of hydrogel nanofibers. Due to hydrophobic effect, intrafibrillar encapsulation prolonged the release of small molecule drugs from MDP hydrogels to achieve long-term drug delivery. For poorly water-soluble small molecule drugs, the “missing tooth” MDP hydrogel provides a tool for localized drug delivery. Together with our previously developed drug delivery strategies, the missing tooth design has expanded the potential of MDP hydrogels to controllably deliver various agents for clinical therapies.19,20,32,33

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ASSOCIATED CONTENT

S Supporting Information *

The Supporting Information is available free of charge on the ACS Publications website at DOI: 10.1021/acs.biomac.6b00309. Solid-phase peptide synthesis protocols; complementary AFM data; MALDI-TOF MS data (SI Figure 1); rheological characterization of SL1 and SL2 hydrogels (SI Figure 2); characterization of SL peptide (SI Figure 3); CD spectra of SL1 hydrogel before and after drug encapsulation (SI Figure 4); negative stained TEM images of SL1 peptide after drug encapsulation (SI Figure 5); degradation curve of diflunisal fluorescence (SI Figure 6); concentration-dependent CD spectra for SL, SL1 and SL2 (SI Figure 7) (PDF)

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AUTHOR INFORMATION

Corresponding Author

*E-mail: [email protected]. Notes

The authors declare the following competing financial interest(s): J.D.H. has stock options in a company which aims to commercialize some aspects of the work described.

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ACKNOWLEDGMENTS The authors appreciate help and support from the Shared Equipment Authority (SEA) at Rice University. The authors would like to thank Carlos A. Origel Marmolejo at Rice University for contribution of Figure 1C,D. The work presented in this manuscript was supported by grants from the Robert A. Welch Foundation (Grant C1557) and from the NIH for J.D.H. (R01 DE021798).

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REFERENCES

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DOI: 10.1021/acs.biomac.6b00309 Biomacromolecules XXXX, XXX, XXX−XXX

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DOI: 10.1021/acs.biomac.6b00309 Biomacromolecules XXXX, XXX, XXX−XXX