Low Molecular Weight Supramolecular Hydrogels for Sustained and Localized In Vivo Drug Delivery

Abstract

Supramolecular hydrogels are emerging as next-generation alternatives to synthetic polymers for drug delivery applications. Self-assembling peptides are a promising class of supramolecular gelator for in vivo drug delivery that have been slow to be adopted despite advantages in biocompatibility due to the relatively high cost of producing synthetic peptide hydrogels compared to synthetic polymer gels. Herein we describe the development and use of inexpensive low molecular weight cationic derivatives of phenylalanine (Phe) as injectable hydrogels for in vivo delivery of an anti-inflammatory drug, diclofenac, for pain mitigation in a mouse model. N-Fluorenylmethoxycarbonyl phenylalanine (Fmoc-Phe) derivatives were modified at the carboxylic acid with diaminopropane (DAP) to provide Fmoc-Phe-DAP molecules that spontaneously and rapidly self-assemble in aqueous solutions upon addition of physiologically relevant sodium chloride concentrations to give hydrogels. When self-assembly occurs in the presence of diclofenac, the drug molecule is efficiently encapsulated within the hydrogel network. These hydrogels exhibit robust shear-thinning behavior, mechanical stability, and drug release profiles to enable application as injectable hydrogels for in vivo drug delivery. Delivery of diclofenac in vivo was demonstrated by a localized injection of an Fmoc-F⁵-Phe-DAP/diclofenac hydrogel into the ankle joint of mice with induced ankle injury and associated inflammation-induced pain. Remediation of pain in the ankle joint was observed immediately after initial injection and was sustained for a period of nearly two weeks while diclofenac controls remediated pain for less than one day. This data demonstrates the promise of these supramolecular hydrogels as inexpensive next-generation materials for sustained and localized drug delivery in vivo.

Introduction

Hydrogels are versatile materials for in vivo delivery and sustained release of encapsulated therapeutic molecules.^1–9 In particular, hydrogels that can be dispensed by injection have great promise for the treatment of localized diseases and ailments, including infections and pain from inflammation.^8–12 Ideal hydrogels for in vivo drug delivery must maintain integrity over long time periods at complex fluid/tissue interfaces, must exhibit shear-responsive behavior to enable delivery by injection, and must release therapeutic concentrations of encapsulated drugs over long periods (minimally days or weeks) to reduce dosing requirements.^7,11–13 In addition, ideal drug delivery hydrogels should be non-irritating and non-immunogenic. Polymer-based hydrogels have been widely employed as drug delivery materials,^14,15 but supramolecular hydrogels formed by the spontaneous self-assembly of peptides, proteins, and other biomolecules have garnered recent attention as more biocompatible next-generation alternatives to synthetic polymer materials.^4,12,16–24 The enhanced biocompatibility and ease of formulation of supramolecular hydrogels compared to polymer hydrogels has inspired research efforts to understand self-assembly phenomena leading to gelation and to engineer the emergent properties of supramolecular gels to form ideal drug delivery materials.^25–31

Self-assembling peptides have shown great promise as biomaterials for drug delivery, although the high cost of synthetic peptides relative to polymers has been an impediment to widespread utilization of peptide-based hydrogels for drug delivery.^32,33 As a result, low molecular weight alternatives to longer peptides, including dipeptides or functionalized amino acids, have been explored as supramolecular hydrogel materials.^33–41 While self-assembling peptides have been shown to spontaneously self-assemble into entangled hydrogel networks in a variety of buffered aqueous solutions, low molecular weight dipeptides and amino acids often have reduced aqueous solubility, which introduces the need for organic cosolvents or complex pH adjustment methods to trigger hydrogelation.⁴¹ In addition, low molecular weight hydrogels often have poor mechanical stability and shear-responsive properties, which complicates the use of these types of materials for in vivo applications.⁴²

Phenylalanine (Phe) derivatives are a privileged class of molecule that have been shown to self-assemble into diverse supramolecular structures, sub-structurally similar to amyloid fibrils, which can entangle to form hydrogel networks.^43–48 The investigation of Phe-derived gelators has been motivated by the need to understand the mechanistic basis for self-assembly phenomena and the relationship between self-assembly and hydrogelation in order to identify cost-effective low molecular weight gelators that are functionally equivalent to peptide-derived counterparts.^47,49–52 N-Fluorenylmethoxycarbonyl phenylalanine (Fmoc-Phe) derivatives are promising as supramolecular hydrogels for sustained drug release applications in vivo.⁵³ Adams and coworkers have shown that Fmoc-Phe and Fmoc-tyrosine (Fmoc-Tyr) hydrogels formed by slow acidification of basic solutions of amino acid by hydrolysis of glucono-δ-lactone (GdL) release dye molecules into layered aqueous solutions in vitro.²⁵ Cao and coworkers have also demonstrated in vitro release of salicylic acid from supramolecular Phe-derived hydrogels.⁵⁴ While drug release and biocompatibility of amino acid-based hydrogels has been studied in in vitro tissue culture systems,^39,55–59 the use of these inexpensive hydrogels for in vivo drug delivery is problematic due to low aqueous solubility of the gelator molecules and the often poor mechanical stability of the resulting gel systems.

We recently reported a series of cationic Fmoc-Phe derivatives that form hydrogels spontaneously in aqueous solutions without the need for organic cosolvents (compounds 1-3, Figure 1).⁶⁰ These Fmoc-Phe derivatives have been modified with diaminopropane (DAP) at the carboxylic acid, which dramatically improves water solubility of cationic 1-3 relative to the parent Phe-derivatives across a broad range of pH conditions. Self-assembly and hydrogelation is triggered simply by addition of physiologically relevant concentrations of sodium chloride, which results in charge screening of the cationic functional groups, promoting self-assembly. As previously observed with fluorinated Fmoc-Phe derivatives,^49,61 fluorination of the benzyl side chain of the Fmoc-Phe-DAP derivatives significantly enhances self-assembly/hydrogelation potential. Based on our initial observations, we hypothesized that these cationic Fmoc-Phe-DAP derivatives would have ideal emergent hydrogel properties for drug delivery applications. Herein, we characterize hydrogels derived from compounds 1-3 for shear-responsive character and in vitro release of diclofenac,⁶² a nonsteroidal anti-inflammatory drug, as a model therapeutic and demonstrate the use of these hydrogels as injectable materials for the localized and sustained release of diclofenac for pain mitigation in in vivo mouse models over two weeks from a single injection of hydrogel. This work demonstrates the practicality of these inexpensive supramolecular hydrogels for in vivo delivery of therapeutics.

Figure 1.

Chemical structures of Fmoc-Phe-DAP cationic gelators. Fmoc-Phe-DAP (1), Fmoc-3F-Phe-DAP (2), and Fmoc-F⁵-Phe-DAP (3).

Characterization of Cationic Fmoc-Phenylalanine Derivatives

We first assessed hydrogel formation and stability at high concentrations in order to ascertain mechanical stability of the gel as a function of interfacial contact with other aqueous solutions. Our initial report of hydrogels from compounds 1-3 were formed at or below concentrations of 20 mM of the gelator.⁶⁰ We found that hydrogels formed at these relatively low concentrations of gelator were mechanically unstable to repeated layering with aqueous solutions, resulting in degradation of the gels over several days. This indicates that gels formed at low concentrations are not likely to be stable for extended time periods after in vivo injection. Thus, we examined hydrogel formation at higher concentrations in order to determine if these hydrogels would be more mechanically stable.

We found that hydrogels formed at concentrations of gelator greater than 30 mM were stable over extended periods of time under conditions of interfacial interaction with other aqueous solutions. To prepare the hydrogels, gelators 1, 2, or 3, synthesized via previously described methods,⁶⁰ were suspended in water and sonicated to provide a finely divided suspension of each gelator in water (see Figure 2A, 2B and Figure S1 in Supporting Information for the appearance of these suspensions). These suspensions were then heated to 80 °C until the gelators were completely dissolved (Figure 2C and D for gelator 1; other gelators shown in Figure S1). Concentrated sodium chloride (NaCl) was added to produce a final concentration of 33.7 mM gelator with 114 mM NaCl in a total volume of 1 mL. After addition of salt, the solution was rapidly mixed (1-2 seconds), and hydrogel formation occurred immediately (see hydrogels of gelator 1 in Figure 2E and F; other hydrogels shown in Figure S1). Detailed protocols for hydrogel formation can be found in the Supporting Information.

Figure 2.

Images illustrating the gelation process of gelator 1 (33.7 mM gelator). Panels A and B show the solution after sonication of the gelator to form a fine suspension. Panels C and D show the dissolution of the gelator after heating to 80 °C. Panels E and F show the hydrogel formed immediately after addition of NaCl and brief mixing by vortex. Hydrogelation occurs immediately after addition of NaCl as indicated by the vial inversion test shown in panel F.

These hydrogels were incubated at 37 °C to simulate physiological temperatures and the mechanical stability of each gel was tested over two weeks to determine suitability for long-term release of therapeutics. These tests were performed by layering saline solution (114 mM NaCl) and phosphate buffered saline solution over the gels and monitoring degradation of the gels at 37 °C as a function of time. Gels formed at gelator concentrations lower than 20 mM showed precipitation of the hydrogel network and reductions in volume over several days due to gel dissolution over time. Gels formed at concentrations of gelator greater than 30 mM, however, were stable for several weeks, depending on the gelator. Within two weeks, Fmoc-Phe-DAP (1) hydrogels showed significant precipitation of the hydrogel network and the gel integrity had been largely compromised (see Figure S2, Supporting Information). Fmoc-3F-Phe-DAP (2) hydrogels became slightly more opaque, but the gels were still intact (Figure S2). Fmoc-F⁵-Phe-DAP (3) hydrogels remained completely intact (Figure S2). The stability of the hydrogels decreases in the order 3 > 2 > 1, indicating that hydrogels of 2 and 3 have the greatest potential as drug delivery vectors for sustained release.

The morphology of each hydrogel network was characterized by transmission electron microscopy (TEM) imaging to understand hydrogel structure and stability. Again, hydrogels were incubated at 37 °C and aliquots were removed after 24 hours and 72 hours for analysis by TEM (detailed protocols in Materials and Methods in Supporting Information). These time points were chosen in order to confirm that the morphology of the hydrogel fiber networks is consistent with previous observations for these gelators at lower concentrations.⁶⁰ Representative images of the self-assembled structures for compounds 1, 2, and 3 after both 24 and 72 hours of incubation can be seen in Figure 3. The morphology of the self-assembled hydrogel networks at high concentrations was similar to those formed at lower concentrations as described in our initial report of these gelators.⁶⁰ Compound 1 self-assembled into flat, tape-like structures that converge and can be seen wrapping up to form nanotubes that are 300–500 nm in diameter (Figure 3A). After 72 hours, larger nanotubes between 500 nm and 1 μm in diameter are the dominant supramolecular structures in hydrogels of compound 1 (Figure 3B and Figure S3, Supporting Information). The addition of a fluorine in the meta position of the phenylalanine ring in gelator 2 slows the progression of fibrils/tapes to tubes. Hydrogels of 2 after 24 hours are composed of thinner, tightly wound fibrils or tapes; within 72 hours, these structures also transform into nanotubes, although a higher density of thinner fibrils are observed compared to compound 1 (Figure 3C and D and Figure S4, Supporting Information). Compound 3, with a fully fluorinated benzyl ring, formed hydrogels composed of twisted fibers and tapes that are much thinner than were observed for either of the other two gelators (approximately 10-20 nm in width, Figure 3E and F and Figure S5, Supporting Information). After 72 hours, nanotubes are also observed (Figure S5), but the ratio of fiber to nanotube is still primarily dominated by thinner fibers.

Figure 3.

Representative TEM images of 33.7 mM hydrogels made from gelator 1 after (A) 24 hours and (B) 72 hours; gelator 2 after (C) 24 hours and (D) 72 hours; gelator 3 after (E) 24 hours and (F) 72 hours; and 1:1 coassembly of gelators 1:3 after (G) 24 hours and (H) 72 hours.

Interestingly, as the degree of fluorination (and hydrophobicity) of the gelator increases, the ratio of fibers/tapes to nanotubes decreases. We hypothesized that a higher ratio of fibers/tapes to tubes would give a more efficiently entangled hydrogel network. To see if a hybrid mix of gelators would alter the morphology of assembled structures, we examined the supramolecular morphology of a coassembled hydrogel of a 1:1 ratio of Fmoc-Phe-DAP (1):Fmoc-F⁵-Phe-DAP (3) (33.7 mM total gelator). Gelator 1 has a high density of nanotubes after 72 hours, while gelator 3 has a low density of tubes and higher density of fibers. There are numerous examples in literature of coassembled peptidic supramolecular structures displaying different emergent properties compared to those of their individual parent peptides and we surmised that the same would be true for mixtures of compounds 1 and 3.^63–65 We found that coassembly of 1 and 3 did indeed produce a hydrogel that was morphologically intermediary compared to the hydrogels of the individual components (Figure 3G and H and Figure S6, Supporting Information). No mature nanotubes were apparent within 72 hours, but the early stages of nanotube formation are apparent through the appearance of flat, twisted ribbons intertwining and folding with a wider helical pitch compared to gelator 1 or gelator 3 alone. This suggests the possibility that emergent hydrogel viscoelasticity can be further tuned by mixing gelators 1-3.

Next the viscoelastic and shear-thinning properties of the hydrogels were characterized via oscillatory rheology. Hydrogels intended for in vivo injection use must undergo shear-thinning under mechanical stress and must be able to reform the gel network after removal of shear forces.^8,11,66 Strain sweep measurements were performed to determine the linear viscoelastic region for each hydrogel (Figure S7). Frequency sweep experiments were then performed on each hydrogel at 1% strain, which is within the linear viscoelastic region for each material, to determine the viscoelasticity of each hydrogel as a function of the storage modulus (G′) and loss modulus (G″) values (Figure 4). Hydrogels of compounds 1, 2, and 3 had G′ and G″ values of 383 ± 100 Pa and 59 ± 17 Pa (1); 21311 ± 2057 Pa and 3973 ± 419 Pa (2); and 10776 ± 902 Pa and 2273 ± 209 Pa (3) respectively. Interestingly, the 1:1 coassembled hydrogel of 1 and 3 displayed a storage modulus of 17109 ± 1925 Pa (loss modulus is 2279 ± 163 Pa), over forty times higher than the hydrogel of compound 1 and nearly twice that of compound 3, highlighting the effectiveness of coassembly for the alteration of emergent properties of hydrogels. In general, hydrogels with higher fiber to nanotube ratios (hydrogels 2, 3, and 1:3) had significantly greater elastic character than hydrogels of compounds with higher nanotube morphology (hydrogel 1). Since the ratio of fibers to nanotubes in these hydrogels is difficult to precisely quantify, the correlation of fiber/nanotube ratio to emergent viscoelasticity is not perfect, as evidenced by hydrogel 2 having a higher storage modulus than hydrogel 3 even though the latter has an apparent lower density of tubes in the TEM images shown. It is possible that other effects, including differences in dipolar and quadrupolar charge density between gelators 2 and 3 also influence the viscoelasticity of these hydrogels.^47,61

Figure 4.

Representative frequency sweep oscillatory rheology measurements of Fmoc-Phe-DAP hydrogels (33.7 mM). G′ and G″ values (Pa) are represented with closed and open circles, respectively. Frequency sweep for gelator 1 is represented in green, gelator 2 in red, gelator 3 in blue, and a 1:1 coassembly of gelator 1:3 in cyan.

The shear-recovery character of each hydrogel was then characterized using dynamic time sweep experiments in which viscoelastic parameters were monitored as alternating conditions of low strain (1%) and high strain (100%) were applied to the hydrogel (Figure 5). Under high strain conditions, the gel network is disrupted, resulting in loss of hydrogel character (shear-thinning). Gels that exhibit shear-recovery behavior show a return of hydrogel viscoelasticity upon a return to low strain conditions. All four hydrogels exhibit shear-thinning properties indicated by the “breakage” of hydrogels at high strain (gel-sol transition) followed by nearly immediate recovery of gel rigidity (sol-gel transition) upon a return to low strain (Figure 5). This is evident through a reduction in and inversion of G′ and G″ values when a strain of 100% was applied between 300-500 and 750-950 seconds followed by immediate recovery of the initial G′ and G″ values after removal of high strain. This cycle was repeated through two iterations of high strain, and reformation was nearly instantaneous for all gels except for the Fmoc-Phe-DAP (1) hydrogels, the weakest of the hydrogels, where reformation of the gel took over 1 minute as evident by the gradual increase of the G′ value after removal of high strain. The shear-recovery properties of these cationic hydrogels are ideal for in vivo injection.

Figure 5.

Rheological shear-recovery data for 33.7 mM hydrogels comprised of (A) gelator 1, (B) gelator 2, (C) gelator 3, and (D) 1:1 coassembly of gelators 1:3. G′ and G″ (Pa) are indicated with closed and open circles, respectively.

Drug Release Profiles from Cationic Fmoc-Phenylalanine Derivatives

We next characterized the in vitro drug release profiles for all four hydrogels to compare the sustained release of molecules from each material. Release assays were conducted using the nonsteroidal anti-inflammatory drug diclofenac as a model compound. Diclofenac loaded hydrogels were prepared by first dissolving diclofenac in water, followed by addition of the gelator and solubilization by sonication and heating. Gelation was triggered by addition of sodium chloride solution (see Materials and Methods in Supporting Information for the detailed protocol). The hydrogels had a final concentration of 33.7 mM gelator and contained 5 mg mL⁻¹ (15.7 mM) diclofenac to ensure a high enough effective drug concentration for eventual pain mitigation in vivo. The release of diclofenac into an interfacial solution of phosphate buffered saline (PBS, pH 7) was carried out by adding 4 mL of PBS solution above the gel and incubating the biphasic mixture in a sealed vessel at 37 °C. Aliquots of the PBS were removed over time and the concentration of diclofenac released was determined for each aliquot at each time point. The ratio of the amount of diclofenac released after t minutes have elapsed to the total amount of diclofenac loaded in the hydrogel (M_t/M_∞) was plotted against time (t, min) (Figure 6A) (see Materials and Methods, Supporting Information for details).^25,67 Differences in the saturating concentration of diclofenac release into the reservoir solution can be seen, with the lowest level of release coming from the hydrogel of Fmoc-F⁵-Phe-DAP (3). A second plot was constructed by plotting M_t/M_∞ against t^1/2 (min^1/2) from the initial linear section of the first plot (comprising approximately the first 240 minutes of the release study) (Figure 6B). The diffusion constant, D (m² min⁻¹), was determined by measuring the slope of M_t/M_∞ against t^1/2 in this second plot and setting this value equal to the coefficient of t^1/2 (min^1/2) above in order to solve for the value of D (m² min⁻¹) using the relationship shown in the non-steady state diffusion model described in equation 1 (see Materials and Methods in Supporting Information).⁶⁷

Figure 6.

(A) Diclofenac release profiles indicating the ratio of drug molecules released (M_t/M_∞) vs. time in minutes from 33.7 mM hydrogels of 1 (green), 2 (red), 3 (blue), and 1:1 coassembly of 1:3 (cyan) all containing 5 mg/mL diclofenac. (B) Linear region of diclofenac release indicated by M_t/M_∞ vs t^1/2 used to calculate diffusion coefficients equal to 9.54 × 10⁻¹³ m² min⁻¹ for diclofenac release from hydrogel 1 (green), 9.71 × 10⁻¹³ m² min⁻¹ from hydrogel 2 (red), 1.23 × 10⁻¹³ m² min⁻¹ from hydrogel 3 (blue) and 1.75 × 10⁻¹² m² min⁻¹ from a hydrogel of 1:1 coassembly of 1:3 (cyan).

The in vitro release profiles of diclofenac from all four hydrogels were found to be similar, with minor differences in diffusion constant. The release rate of diclofenac from the 33.7 mM hydrogels under these conditions was found to follow the trend 2 ≥ 1 > 3 (9.71 × 10⁻¹³ m² min⁻¹, 9.54 × 10⁻¹³ m² min⁻¹, and 1.23 × 10⁻¹³ m² min⁻¹ respectively). The release rate of diclofenac was greatest from the hybrid hydrogel comprised of 1:1 coassembly of gelators 1:3 and is equal to 1.75 × 10⁻¹² m² min⁻¹. In general, the rate of release of diclofenac increased almost linearly over approximately the first 8 hours followed by a slow decrease in rate of release until the diclofenac release reached a saturating concentration over the next two days. The values of diffusion constant from each of the gels makes it difficult to draw meaningful conclusions about how the moderate differences may be linked to differences in network morphology or hydrogel viscoelasticity. For in vivo applications it is sufficient to understand that the difference in release rates of diclofenac from the various hydrogel formulations are exceedingly minor.

The total amount of diclofenac released from the hydrogels followed the trend 1:3 > 1 > 2 > 3. This trend also does not strictly follow the trends observed for hydrogel viscoelasticity or network morphology (ratio of tube structures to fibers/tapes), indicating that differences in the total amount of diclofenac released cannot be explained entirely by these emergent properties of the hydrogels. A recent study of release of diclofenac from positively charged phosphonium gels indicated that the diclofenac release profiles were unaffected by changes in pH but diclofenac release was slower from hydrogels containing triphenylphosphine compared to hydrogels with non-aromatic phosphines.⁶⁸ Presumably, aromatic diclofenac molecules participate in specific π-π interactions with aromatic triphenylphosphine that is not possible with non-aromatic phosphines. This suggests that the molecular structure of the gelators in this study may play a role in the differences in the amount of diclofenac released at saturation. Fmoc-F₅-Phe-DAP (3) has an altered quadrupole in the aromatic benzyl side chain that may result in more attractive π-π interactions between the gelator and cargo for this gelator. This would be observed to a lesser degree in the monofluorinated Fmoc-3F-Phe-DAP gelator (2) and the nonfluorinated Fmoc-Phe-DAP gelator (1), which matches the general trend as evidenced by the varying release from hydrogels comprised of 1, 2, or 3.

In Vivo Sustained Release of Diclofenac

Finally, we explored validation of these Fmoc-Phe-DAP derived hydrogels for the in vivo delivery of diclofenac for the functional relief of pain in a mouse model. For this study we deemed it unnecessary to test all four hydrogel formulations. We eliminated hydrogels of Fmoc-Phe-DAP (1) as a candidate for in vivo delivery due to the lower mechanical stability of these gels. The gels of 2, 3, and 1:3 were similar in terms of most emergent properties tested in vitro. Ultimately, we chose to use Fmoc-F⁵-Phe-DAP (3) for in vivo analysis on the basis of the robust mechanical stability and the slower rate of release and lower level of saturating concentration of diclofenac from these gels. Based on these properties, we reasoned that Fmoc-F⁵-Phe-DAP (3) hydrogels would be the strongest candidate to give sustained release of diclofenac in vivo.

In vivo validation of Fmoc-F⁵-Phe-DAP (3) hydrogels for functional delivery of diclofenac was carried out using an induced pain model in mice.⁶⁹ Acute inflammatory pain was induced in mice by an intra-articular administration of complete Freund’s adjuvant (CFA) into a hind limb ankle joint, resulting in pronounced hind paw sensitivity. Two days after pain induction, diclofenac solution (0.1 mg/ml in physiological saline, 10 μl), diclofenac in Fmoc-F₅-Phe-DAP (3) hydrogel (5 mg/ml, 10 μl), or vehicle (H₂O) in Fmoc-F₅-Phe-DAP (3) hydrogel (10 μl, formulated as described previously) was administered by injection to the afflicted ankle joint of various animal groups under anesthesia (1% isoflurane). A fourth animal group did not receive CFA administration (no pain induction), and Fmoc-F₅-Phe-DAP (3) hydrogel (10 μl) was administered into a hind limb ankle joint to determine if the gel itself resulted in irritation or inflammation. The effective concentration of the direct diclofenac control injection (no gel) was an estimate based on the percentage of diclofenac released from a 1 mL hydrogel comprised of gelator 3 over 72 hours with a measured diffusion coefficient of 1.76 × 10⁻¹¹ m² min⁻¹ via in vitro analysis. Hind paw sensitivity was monitored in the affected leg of all animal groups over 14 days. The circumference of the affected ankle in all test/control groups was also monitored in order to ascertain any inflammation due to the gel itself.

Within hours of diclofenac, diclofenac gel, or control injections, pain as monitored by relative percentage of hind paw sensitivity decreased significantly in both groups that received the diclofenac gel (Dcf/F⁵-Phe) or solution (Dcf) injections as compared to gel lacking diclofenac (F⁵-Phe) group in which paw sensitivity did not decrease (Figure 7A and B). Only 24 hours after injection however, the mechanical sensitivity in the group administered Dcf solution returned and leveled off near 80% over the final 13 days. The Dcf/F⁵-Phe hydrogel group, however, had a 50% reduction in sensitivity within 12 hours of hydrogel injection until day 11, where sensitivity then gradually increased between days 11-14 to the same level as Dcf solution (70%). Paw sensitivity in the hydrogel alone group (F⁵-Phe) remained at 90% throughout the experiment, essentially unchanged from the untreated group. Comparatively, the control group that was not subjected to CFA-induced pain that was injected with equal volumes of F⁵-Phe hydrogel displayed between 20-40% sensitivity throughout the course of the experiment indicating an insignificant response to the presence of the hydrogel alone (Figure S10, Supporting Information).

Figure 7.

(A) Mechanical sensitivity measured over time for groups of 6 mice. One group treated with vehicle F⁵-Phe alone (white squares), one with 10 μL of Dcf/F⁵-Phe (3) containing 5 mg/mL diclofenac (black squares), and one with 10 μL of a 0.1 mg/mL Dcf solution (grey circles), the estimated effective concentration of diclofenac based on in vitro release profiles. The percent sensitivities are all relative to 100% sensitivity that was established by the group of mice who received CFA-induced pain and were only treated with F⁵-Phe (3) hydrogel alone. (B) Change in mechanical sensitivity over time for all groups of mice. * indicates a significant difference with P < 0.05 and ** indicates a statistical difference with P < 0.01. Striped F⁵-Phe (3) group indicates the non-CFA induced pain group treated with F⁵-Phe (3) hydrogel alone. (C) Graph of the ankle circumference of all four groups of mice with statistical differences indicated.

These results clearly demonstrate that the Fmoc-F⁵-Phe-DAP (3) gelator is a highly effective injectable hydrogel for the localized in vivo functional delivery of diclofenac. These diclofenac hydrogel formulations also show sustained drug release for nearly two weeks post injection for effective pain mitigation. Ankle circumference of the affected limbs was measured for all groups and at day five, the circumference of the ankles of the non CFA-induced mice treated with the F⁵-Phe gel had significantly reduced circumferences than the CFA-induced mice treated with F⁵-Phe (P < 0.01) indicating the affliction is not due exclusively to the presence of hydrogel (Figure 7C). Compared to the CFA-induced group that received the F⁵-Phe gel, slightly lower circumferences were observed at day five in both Dcf/F⁵-Phe or Dcf solution groups. After 14 days, the circumference of the ankles in the Dcf/F⁵-Phe group were significantly lower than the F⁵-Phe group (P < 0.01), and slightly lower than the Dcf solution group, but within error of this result. Thus, the hydrogel formulations themselves do not cause measurable discomfort as determined by hind limb sensitivity. Together these data confirm that cationic Fmoc-Phe-DAP derived supramolecular hydrogels are viable injectable materials for localized, sustained drug release in vivo.

Finally, we conducted a study to confirm that the hydrogel remains intact in vivo over the course of these studies. Fluorescein was added to Fmoc-F⁵-Phe-DAP (3) hydrogels as a fluorescent reporter and these hydrogels were injected subcutaneously into the hind limb of mice. After 24 hours or 10 days, the mice were sacrificed and the hind limb was immediately frozen. The hind limb was then sliced using a cryostat and fluorescence imaging was used to confirm the location of the hydrogel. It was found that these hydrogels remained intact after 10 days (Figure S11, Supporting Information). Thus, the mitigation of pain over time from diclofenac hydrogels is consistent with sustained release of the drug from the hydrogel and not from deterioration of the hydrogel over the course of the experiment.

Conclusion

Herein we have described the development of low molecular weight supramolecular hydrogels that have been applied to in vivo localized and sustained drug delivery. Low molecular weight supramolecular hydrogels are advantageous to synthetic polymer hydrogels and to peptide/protein supramolecular hydrogels in terms of biocompatibility, cost, and avoidance of laborious synthesis and formulation protocols. Heretofore, the emergent molecular properties of existing low molecular weight self-assembling gelators, including low aqueous solubility and stability, lack of shear-responsive characteristics, and incompatibility with biological solvents and conditions, have impeded the use of these materials for widespread in vivo drug delivery. In this work we have surmounted these obstacles by modifying the privileged self-assembly motif of Fmoc-Phe derivatives with cationic groups that enhance water solubility and enable facile and rapid self-assembly in aqueous solutions simply by the addition of physiologically appropriate concentrations of sodium chloride. This simple formulation method is ideal for the facile encapsulation of small molecule drugs within the hydrogel network, as was demonstrated by the inclusion of the anti-inflammatory diclofenac in the hydrogels. Further, we demonstrated that these hydrogels are shear-responsive and have ideal viscoelastic character for delivery by injection and to maintain integrity at the interface of biological tissues and fluids over periods of weeks. Finally, these hydrogels release drugs over periods of weeks, as demonstrated by in vivo diclofenac delivery for functional pain remediation in localized injury-induced inflammation. Based on these properties, these materials have great potential to fulfill the promise of low molecular weight supramolecular hydrogels as next-generation drug delivery vectors.