Self-Nanoemulsifying Electrospun Fiber Enhancing Drug Permeation

Research Article www.acsami.org

Cite This: ACS Appl. Mater. Interfaces XXXX, XXX, XXX−XXX

Self-Nanoemulsifying Electrospun Fiber Enhancing Drug Permeation Yi Xiang,†,‡ Jing Liang,† Lili Liu,‡ Fei Wang,† Lianfu Deng,† and Wenguo Cui*,† †

Shanghai Key Laboratory for Prevention and Treatment of Bone and Joint Diseases, Shanghai Institute of Traumatology and Orthopaedics, Ruijin Hospital, Shanghai Jiao Tong University School of Medicine, 197 Ruijin 2nd Road, Shanghai 200025, P. R. China ‡ Orthopedic Institute, Soochow University, 708 Renmin Road, Suzhou, Jiangsu 215006, China

ACS Appl. Mater. Interfaces Downloaded from pubs.acs.org by MACQUARIE UNIV on 02/18/19. For personal use only.

S Supporting Information *

ABSTRACT: Electrospun fibers are excellent drug carriers and tissue engineering scaffolds. However, approaches to promote drug permeation in tissues with such carriers remain of great interest. Here, we propose a Quality-by-Design strategy to enhance drug permeation with self-nanoemulsifying electrospun fibers. Owing to the nanoemulsion which formed spontaneously when the polymer contacts aqueous solution such as body fluid, the resulting drug-laden fibrous membrane exhibits an outstanding drug permeation and therapeutic enhancement effect in a Franz cell experiment with ex vivo abdomen skin of rats, an artificial connective tissue model, and an in vivo rheumatoid arthritis model in rats. Meanwhile, the material also shows the capacity of rational regulation on the rate of drug release. These features of the present strategy establish our material as a new efficient approach for various clinical conditions calling for cure. KEYWORDS: electrospinning, self-nanoemulsifying drug delivery system, nanoemulsion, topical drug delivery, drug permeation

■

molecules.22,23 Chemical methods boost the permeation of drugs by enhancing their passive diffusion taking advantage of the partial fluidizing of the cellular barrier caused by chemical agents such as alcohols, fatty acids, Azones, and so forth,24 whereas the physical methods, on the other hand, infuse drugs by the aid of various external energies through different pathways, mature approaches including iontophoresis, sonophoresis, electroporation, and microneedles.25 However, both methods inevitably cause undesirable irritation and injury of the barrier,22 which is acceptable when applied on healthy and unbroken skin, but not in the case of other sites of body. Meanwhile, physical methods often involve a complicated process in manufacturing as well as application,23 adding to the difficulty of their co-operation with electrospun fibrous membranes. Methods based on nanoencapsulation26 seem to be the most promising strategy to enhance drug permeation with electrospun fibrous membranes. Prevailing nanoencapsulation systems include nanoparticles, liposomes,27 micelles, dendrimers, and nanoemulsions (NEs). These nanocarriers promote the permeation of drugs by carrying them through the intercellular and follicular route taking advantage of their size and composition. There have also been approaches to modify the nanocarriers with receptor-specific ligands to enhance the drug permeation through an intracellular route.28 However, most of these carriers themselves suffer from drawbacks, having

INTRODUCTION Electrospun fibrous membranes have been under extensive research as excellent drug delivery systems (DDS) and tissue engineering scaffolds.1,2 These membranes are of special significance as drug carriers for local delivery of therapeutic molecules in virtue of their huge aspect ratio,3 fibrous microstructure,4 as well as easiness in fabrication. In this context, earnest endeavors have been devoted to the expedition of the potential of electrospun fibrous membranes to exhibit advanced properties as DDS. So far, spinning methods such as blending,5,6 co-axial,7,8 emulsion,9,10 nanoparticle preloading,11,12 micro-sol,13 and micro-gel14 have been developed to incorporate hydrophilic,6 lipophilic,10 and insoluble drugs15 into the fibers and realize their release in burst,16,17 sustained,18 and responsive17,19 manners. It remains great challenge, nevertheless, that the therapeutic outcomes are still constrained by the unsatisfactory absorption and permeation of these drugs in the target tissues or lesions, because drugs tend to be deposited near the surface where the membranes are applied instead of diffusing to the depth of the tissue, especially for the case of drug delivery in solid tumors which often form a dense matrix20 and transdermal therapy for noncutaneous medical conditions such as joint and muscle diseases.21 Hence, there is an essential demand to develop an electrospun fibrous membrane-based DDS that enhances the permeation of drugs in tissues. Three categories of strategies have been developed to enhance drug permeation: chemical methods, physical methods, and methods based on encapsulation of the © XXXX American Chemical Society

Received: December 20, 2018 Accepted: January 30, 2019

A

DOI: 10.1021/acsami.8b21967 ACS Appl. Mater. Interfaces XXXX, XXX, XXX−XXX

Research Article

ACS Applied Materials & Interfaces Scheme 1. Illustration of the Process of the Drug Absorption Enhancement Induced by an SNE-FM

designed to occur spontaneously, the qualities of this preparation such as content uniformity, dissolution, and other properties are easily controlled, which meets the principle of Quality-by-Design (QbD) and has a high potential to be translated to market.

difficulty in meeting the basic requirements of an ideal drug permeation enhancing DDS at the same time, including good stability, high drug loading efficiency (LE), low selectivity toward drugs and tissues, manufacturing and applying process without high-demanding techniques, and easiness in quality control. Although NE being an exception in terms of these drawbacks,29,30 there is also a demand of compatibility of the carrier with the electrospinning process or with the fibrous nanostructure to be incorporated into electrospun fibrous membranes. So far, it has never been realized to deliver nanoemulsions with electrospun fibrous membranes. Here, we propose a general strategy to prepare a novel selfnanoemulsifying electrospun fibrous membrane (SNE-FM) that forms autonomously and enhances drug permeation spontaneously, for which we name this system as an initiative drug delivery system (iDDS). By co-dissolving the selfnanoemulsifying drug delivery system (SNEDDS) and polymers in an optimized alcohol solvent, all the components showed supreme affinity toward each other and assembled into a homogeneous system which was named as a self-nanoemulsifying polymer (SNE-P). Thanks to its resistance to external forces, the system could be further fabricated into different preparations arbitrarily, including SNE-FM. When contacted with water, the SNEDDS components exhibit equivalent self-nanoemulsifying capabilities with the normal SNEDDS that spontaneously assemble into NEs to enhance drug permeation in various tissues effectively. Furthermore, since both the fabrication and the drug delivery processes are

■

RESULTS AND DISCUSSION The fabrication and physicochemical properties of SNE-FM. SNE-Ps were successfully produced with the SNEDDS and polymers by the aid of the hydroxyls of the optimized alcohol solvent. Assemblies of SNE-P with different SNE degrees were accomplished by adjusting the concentration of each component as presented in Table S1. As soon as the assembling was completed, a viscoelastic solution was formed, on which a high-voltage electric field was subsequently applied to stretch its droplets into nanofibers. The resulting fibers were able to be nanoemulsified spontaneously in virtue of the existence of SNEDDS in the polymer chain.31 A rotating cylinder covered with aluminum foil was grounded to collect these fibers, eventually making up the SNE-FM. The samples were named after their formulation. For example, a SNE-FM of which the matrix contained 67% (w/w) poly-L-lactic acid (PLLA) and 33% (w/w) SNEDDS and was laden with 10% (w/w) drug of its total weight was denoted as L67@S33@ D10. Once the SNE components contacted with aqueous liquid, a stronger force (hydrogen bonding and hydrophobic force) was produced and the SNE components were disassociated. Subsequently, they self-assembled into lamellar B


Research Article

ACS Applied Materials & Interfaces

Figure 1. Chemical and physical changes brought by the composition of SNEDDS into the FM based on hydrophobic polymer PLLA. (a) Fourier transform infrared spectroscopy data. (b) Water contact angle (WCA) results. (c) Stress−strain curves. (d) Elastic modulus data.

liquid crystals,32 accompanied by a sharp decline in temporary interfacial tension. NE was spontaneously formed under the extremely low interfacial tension when aqueous liquid continued to be added to the lamellar liquid crystal.33 Simultaneously, a hydrophobic drug was confined within the oil nanodroplets or interfacial membrane of the NE, and thus transported to the tissue to be absorbed (Scheme 1). The method to fabricate an SNE-FM was successfully replicated with hydrophobic polymer PLLA and water-soluble polymer poly(vinylpyrrolidone) (PVP). The assembly was investigated in terms of the changes taking place in the FMs physically and chemically. The Fourier transform infrared (FTIR) spectrum of L67@S33@D10 compared to L100@ S0@D10 is presented in Figure 1a. The co-existence of PLLA and SNEDDS generated a new C−O−C stretching vibration, characterized by the peak at 1140 cm−1 in the L67@S33@D10 spectrum. On the other hand, a red shift and hyperchromicity took place in the L67@S33@D10 spectrum at 1192 cm−1, the typical C−O antisymmetric stretching vibration peak of PLLA,34 compared to that of L100@D0@D10 (1182 cm−1), indicating that the regularity of the molecule arrangement within the system was reduced when the SNEDDS was present, which means that the SNEDDS was uniformly embedded in the polymer. Hyperchromicity also took place at 1762 cm−1, the characteristic peak of PLLA CO stretching vibration, and in the FTIR spectrum of the PVPbased SNE-FM (Figure S1). To conclude, driven by the high compatibility between the polymer and SNEDDS components, the SNEDDS was embedded in the polymer chains to form a uniform FM which was stabilized by their high affinity. As shown in Figure 1b, the water contact angle (WCA) was up to 114.32° in the drug-laden regular PLLA FM, indicating its hydrophobicity. On the other hand, the other three FMs incorporated with the SNEDDS absorbed the water droplet immediately when contacted with water, despite the SNEDDS/PLLA proportion. It could be deduced that the SNEDDS was assembled into the PLLA chain and enhance the randomness of the latter, resulting in a hydrophilic SNE-P.

This was in accordance with the FTIR results. Unfortunately, the WCA experiment was not conducted on PVP-based FMs because the PVP matrix itself was water soluble. The stress− strain curves of the PLLA FM composited with different proportions of SNEDDS are presented in Figure 1c. It was obvious that the contrails of L100@S0@D10, L67@S33@ D10, and L50@S50@D10 curves were close, whereas the stress of L90@S10@D10 was significantly higher than that of the other three samples at the same strain. This was likely to be contributed by the extra force produced by the high affinity between the SNEDDS and the polymer. When the SNEDDS was incorporated into PLLA at the proportion of 10/90, the resulting FM showed a notable increase in the tensile strength. However, when SNEDDS/PLLA proportions continued to rise to 33/67 and 50/50, their stress−strain curves regressed to the curve of the regular PLLA FM. This was probably because the SNEDDS itself was a semisolid that did not contribute to the mechanical strength, therefore the mechanical enhancement in L90@S10@D10 results entirely from the stabilizing force between the SNEDDS and the polymer. On the other hand, the tensile strength of the FM showed a decrease when the SNEDDS proportion reached a certain level for the same reason. Interestingly, although L90@S10@D10 exhibited a distinct performance compared to other samples in tensile strength, its elastic modulus showed no difference compared to that of L100@S0@D10. L67@S33@D10 had almost same elastic modulus as L90@S10@D10, whereas the elastic modulus of L50@S50@D10 was significantly lower than that of L100@S0@D10 (P < 0.05) (Figure 1d). Despite the nondifference in the data of L90@S10@D10 and L67@S33@ D10 from L100@S0@D10, the elastic modulus reduction effect was detectable when comparing the results of all the four groups. It was not until the incorporation proportion reached a certain level that the elastic modulus showed a significant difference. These results suggested that the incorporation of the SNEDDS into the PLLA polymer not only rationally regulated its drug release mechanism and enabled it to enhance drug permeation (as demonstrated later), but also retained the C


Research Article


Figure 2. Microstructure of the SNE-FMs under SEM and their molecular model. (a) L100@S0@D10, (b) L90@S10@D10, (c) L67@S33@D10, (d) L50@S50@D10, (e) V100@S0@D10, (f) V90@S10@D10, (g) V67@S33@D10, (h) V50@S50@D10, (i) scheme of the bottom-up model of the SNE-FM.

reducing the environmental relative humidity (RH) of electrospinning and preservation could effectively mitigate the adhesion between fibers. When observed with SEM immediately after electrospun at RH under 35%, adhesion barely appeared to V50@S50@D10 (Figure S2). Therefore, it was inferred that the incorporation of the SNEDDS with water-soluble polymers did not affect the fiber formation during electrospinning, but caused moderate melting of the fibers after fabrication. The mechanism of the melting was possibly due to the prevention of solvent volatilization prevention and water absorption enhancement in the fiber caused by the SNEDDS but required further study to be determined. The smooth and homogeneous micromorphology of the SNE-FMs was in accordance with FTIR, WCA, and mechanical test results. Combining these understandings, we propose a bottom-up model to show the molecule arrangement in this system (Figure 2i). SNE Process Study. The SNE function means that the SNEDDS can nanoemulsify itself to form NE without or with a minimum external force when contacted with an aqueous fluid. It was well-established that the material went through three stages during the whole process, that is, the SNEDDS, lamellar liquid crystal and NE,33 and the intermediate liquid crystal had a different electron density distribution from that of the SNEDDS and NE. Therefore, small-angle X-ray scattering

original mechanical performance of PLLA, making it an ideal drug carrier as well as potential tissue engineering scaffolds. Morphology and Bottom-Up Model of the SNE-FM. The micromorphology of the SNE-FMs was observed by a scanning electronic microscope (SEM) and is presented in Figure 2. Figure 2a is the picture from the single component PLLA FM (L100@S0@D10), whereas Figure 2b−d are the pictures from PLLA/SNEDDS proportion 90/10 (L90@S10@ D10), 67/33 (L67@S33@D10), and 50/50 (L50@S50@ D10), respectively. The FMs were fabricated under the same electrospinning conditions and prepared with identical pretreatment procedures. Compared to Figure 2a, no visible difference was observed in Figure 2b−d. All the four FMs were smooth and homogeneous on the surface, which suggested that the SNEDDS was uniformly inlaid in the PLLA chain on the molecular level. The arrangement and diameter of the fibers did not show noticeable variation, indicating that the existence of the SNEDDS did not affect the fiber formation during the electrospinning process. Figure 2e−h were pictures from PVPbased FMs, V100@S0@D10, V90@S10@D10, V67@S33@ D10, and V50@S50@D10, respectively. The diameter, arrangement and the texture of V100@S0@D10 and V90@ S10@D10 were consistent, whereas the adhesion between fibers occurred to V67@D33@D10 and became even severe on V50@S50@D10. Further experiments showed that D


Research Article


Figure 3. Small-angle X-ray scattering analysis of (a) L100@S0@D10, (b) L67@S33@D10, (c) V100@S0@D10, and (d) V67@S33@D10, indicating the SNE process in SNE-FMs when contacted with water. Black curve = dry sample, red curve = moisturized sample, blue curve = leachate/solution of sample. Curves in desaturated colors in each figure: parallel test outcome from SNEDDS samples. Macroview of the leachate/ solution of (e) L100@S0@D10 on the left and L67@S33@D10 on the right, (f) V100@S0@D10 on the left and V67@S33@D10 on the right and (g) ultrapure water (UPW) on the left and SNEDDS on the right.

Figure 4. Dynamic Light Scattering analysis results. Size distribution of the PLLA-based SNE-FM leachate (a) and PVP-based SNE-FM (d), compared with that of NE transferred from SNEDDS, ζ-potential distribution of PLLA-based SNE-FM leachate (b) and PVP-based SNE-FM (e) compared with that of UPW, (c) data from (a) and (b) presented in the histogram, (f) data from (d) and (e) presented in the histogram.

(SAXS) could be utilized to confirm its existence.32 The moisturized FM (which logically contained lamellar liquid crystal) and FM lixivium (which was logically composed of NE) were prepared by adding water to the original FM. L67@ S33@D10 and V67@S33@D10 were analyzed as representative SNE-FMs, and L100@S0@D10 and V100@S0@D10 were pre-treated and analyzed parallelly as their control,

respectively (Figure 3a−d). The same procedures were conducted on the regular SNEDDS parallelly to collect the positive control data. Comparing Figure 3a to 3b, L100@S0@ D10 and L67@S33@D10 exhibited an identical intensity/ scattering curve originally, showing that the assembled SNEDDS in the chain did not affect the electron density distribution of the whole system. When wetted, on the other E


Research Article


Figure 5. Accumulated penetration percentage-time curve of the drug through the dialysis membrane from (a) PLLA-based SNE-FMs varying by the SNE degree, (b) PLLA-based SNE-FMs varying by drug loading, (c) PVP-based SNE-FMs varying by the SNE degree, V50@S50@D10, (d) PVP-based SNE-FMs varying by drug loading.

hand, the curve of L100@S0@D10 coincided with the original sample, whereas the L67@S33@D10’s curve appeared differently, showing a high intensity through the whole examined process without peaks showing up. Compared to the wetted SNEDDS sample, which existed as a lamellar liquid crystal, L67@S33@D10 did not present the characteristic peak of the lamellar liquid crystal structure as expected, however, it did make a difference on the scattering result in the range of 2−5° around the original X-ray beam. For the case of PVP-based FMs, as Figure 3d,e indicated, despite the fact that the original samples of V100@S0@D10 and V67@S33@D10 had an identical intensity/scattering curve, the wetted V67@S33@ D10 induced a peak at the scattering vector = 0.61 nm−1, whereas wetted V100@S33@D10 did not. Although the position of the peak was not exactly in accordance with that of the characteristic peak of the wetted SNEDDS (0.36 nm−1), it was still reasonable to take the peak as a signal that the initial self-assembly of the SNEDDS took place in the V67@S33@ D10 sample, since PVP could interfere with the SNE process as an amphiphilic polymer. In Figure 3a−d, all the lixivium from the four samples produced identical flat curves, whereas NE induced a wide peak at 0.35 nm−1. Referring to the macroview pictures of the lixivium and NE in Figure 3e−g which showed obvious blue opalescence (macrofeature of nanodispersions) as well as the dynamic light scattering (DLS) data presented, it could be inferred that low density of nanodroplets in the lixivium was the reason why they did not show the same peaks as NE did. To conclude, the SAXS data indicated that the formation of SNE intermediate, lamellar liquid crystal in the FMs was different from that of the regular SNEDDS, but its efficiency as an intermediate of NE formation was still present. This was probably because the polymer matrix which was uniformly embedded and blended with the SNEDDS on the molecular level had an effect on the initial self-assembly.35

To further demonstrate the SNE function of the fabricated SNE-FM, DLS analysis was utilized to characterize the SNEFM lixivium (Figure 4). As Figure 4a shows, the size of regular NE droplets distributed around 28 nm, adjoining it was a 24 nm peak of L67@S33@D10 lixivium, verifying the presence of nanodroplets in the FM lixivium. Meanwhile, L67@S33@D10 lixivium also had a peak at 155 nm, which fell very close to the only peak of L100@S0@D10 lixivium (155 nm), both deduced to be contributed by the minor fibers detached from the FM. Figure 4b showed the ζ-potential distribution of the samples. The ζ-potential of ultrapure water (UPW) was around −7.5 mV, whereas that of NE was around 6 mV, indicating the existence of the acting force between the droplets which made the potential of the system positive and kept the dispersion stable. The ζ-potential of L100@S0@D10 lixivium was around −22 mV, on the left side of that of UPW. This was contributed by the large amount of carboxyl in the PLLA chain dissociated in the lixivium as well as the IBU released from the fibers. The peak of L67@S33@D10 lixivium was on −21 mV, 1 mV larger than that of L100@S0@D10 lixivium, indicating that less negative charges were produced in this lixivium and this was inferred to be attributed to the nanodroplets that wrapped up the IBU. On the other hand, the peak showed up with an obvious trailing on its right side, suggesting that extra positive charges were produced, which was in accordance with the ζ-potential results of NE. To conclude, together with the droplet size distribution, ζpotential analysis authenticated the formation of NE from the self-nanoemulsification of L67@S33@D10, the SNE-FM. Figure 3d,e were size distribution and ζ-potential results of lixivium from the PVP-based FM. Similar to the results from PLLA samples, V100@S0@D10 lixivium did not contain particles of diameters between 0 and 100 nm, whereas showing a particle size distribution at round 700 nm which represented the micelles consisting of amphiphilic PVP molecules. On the F


Research Article

ACS Applied Materials & Interfaces other hand, the ζ-potential distribution of V67@S33@D10 had two peaks at 76 and 620 nm, the former attributed to the self-nanoemulsification outcome of V67@S33@D10 and the latter came from the PVP micelles. The result that the V67@ S33@D10 nanodroplet was much bigger than that of NE was imputed to the effect of PVP on the nanoemulsification process. The ζ-potential results of PVP-based FMs showed coherent transitions with that of PLLA. Since PVP was neutral, the ζ-potential of −11 mV presented by the V100@S0@D10 lixivium was completely attributed to the IBU, whereas the −2 mV presented by V67@S33@D10 suggested that part of the IBU was directly released into the UPW and the rest was encapsulated into the nanodroplets. To sum up, both hydrophobic polymer (PLLA)-based and water-soluble polymer (PVP)-based SNE-FMs fully exhibited their SNE capability to form NEs with aqueous solution, and the nanodroplet thus formed could effectively encapsulate drugs laden in the fibers. SNE-FM as the Drug Delivery System. The drug release profile was investigated with modified Valia-Chine cells, as described in the section, Investigation on in Vitro Drug Release. As illustrated in Figure 5a, L90@S10@D10, L67@ S33@D10, and L50@S50@D10 showed a distinct release profile with that of L100@S0@D10. Thirty minutes after the experiment began, 46.34, 54.25, and 74.83% of the drug laden in the three SNE-FMs were detected in the receptor cell through the model semipermeable membrane, respectively. After another 30 min, the numbers increased to 60.96, 77.02, and 91.83%, respectively. When releasing had been taking place for 4 h, over 99% of the drug was detected in the receptor cell for L50@S50@D10, which exhibited the highest drug release rate because of the highest SNEDDS/PLLA ratio, followed by 97.82% of L67@S33@D10. Due to the relatively low SNEDDS content, the percentage of L90@S10@D10 at the same time point was 76.64% and reached 83.81% after 4 more hours. On the other hand, only 11.16% of the laden drug in L100@S0@D10 had penetrated through the membrane in 1 h and this ratio increased to 25.83% at 8 h. It could be concluded that the release rate of the FM tended to go up with the increase of the SNEDDS/PLLA proportion. This trend was led by two mechanisms. First, the SNE process was promoted by the increase of SNEDDS proportion in the FM matrix, enhancing the extent to which the aqueous solution around the FM was nanoemulsified. Therefore, more nanodroplets were formed to promote the drug dissolution. On the other hand, the higher the SNEDDS/PLLA ratio was, the more hydrophilic the FM was. While maintaining the water-resistant nature of PLLA, the obstruction that the nanodroplets or nonencapsulated drug molecules faced while diffusing was reduced. Hence, it was easier for the drug to be released and to penetrate through the membrane. Although presented in visible form, these data were also analyzed in mathematical models such as the first-order kinetic model, Higuchi model, Weibull model,36 and Ritger−Peppas model37 (eq 1). The data fitted the best with the Ritger− Peppas model (Table S2), which was therefore utilized for further discussion. M t /M∞ = kt n

D10, the variation of k showed a positive correlation with that of SNEDDS/PLLA (S/P). Taking S/P as an independent variable, k as a dependent variable, linear regression revealed that their variation followed the function k = 32.212S/P + 54.218(R2 = 0.9945). However, taking L100@S0@D10 together, the R2 decreased to 0.7114 (k = 60.424S/P + 32.14). It was not perfectly persuasive to reach the conclusion of linearity with only three sample sites, but the high goodness of fit here did provide strong evidence for future prediction of permeation rate of this system with the S/P ratio between 50/ 50 and 90/10. SNE-FMs of which the S/P value was fixed at 33/67 was selected to go through further investigation with a series of drug loading efficiency (LE). As shown in Figure 5b, the release percentage-time curves of the four samples did not show a significant difference, but the different LE of drug in FMs with certain weight resulted in a notable variance in the release rate constant k and accumulate release percentage after 4 h. When the drug occupied 2, 5, and 10% of the total weight of the FM, the k values were 43.11, 57.53, and 71.41, respectively, showing a positive correlation with LE. When the drug proportion continued to increase to 15%, the k value decreased to 62.44 instead of increasing accordingly. Taking LE as the independent variable and k as the dependent variable for linear regression, it could be learned that k = 3.4474LE + 37.782(R2 = 0.9757). Taking L67@S33@D15 together, the R2 decreased to 0.5387 (k = 1.515LE + 46.478). Therefore, when LE is between 2 and 10% (w/w), the release constant k of the PLLA-based SNE-FM could be inferred from the function k = 3.4474LE + 37.782, whereas when the LE exceeds 10%, for example reached 15%, the release rate should be investigated independently. The release profile regulation effects of PVP-based SNEFMs were also investigated with the same method but did not appear (Figure 5c,d), the corresponding Ritger−Peppas model parameters are recorded in Table S2. In Figure 5c, the curves were very close and the difference could hardly be seen in the fitting results in the Ritger−Peppas model, indicating that the existence and the proportion of SNEDDS in the PVP FM produced a very slight effect on the outcome of the drug release quantitatively. In Figure 5d, where S/P was fixed and various LE were investigated, the D2 and D5 groups only produced a release percentage of 60 and 80%, respectively, and drug penetrated the membrane completely when LE reached 10 or 15%. Fitted in the Peppas−Sahlin model38 (as described in Figure S6), there was no difference in the drug release mechanism as well. It was suggested by the DLS data that PVP in the SNE-FM formed micelles of 620 nm in diameter during the SNE process. The micelles could further make up a network to confine the nanodroplets and drug molecules and act as a barrier of their diffusion. When LE was high enough to produce a sufficient concentration difference across the membrane, this barrier is neglectable, but it was highlighted when the concentration difference was not significant enough, such as in the cases of D2 and D5. SNE-FM Enhancing Drug Permeation. NEs can effectively enhance the transdermal and transmucosal permeation of drugs to promote their absorption and activity39−41 and as demonstrated, the SNE-FM can produce NEs spontaneously when contacted with aqueous solution. Therefore, we further investigated the ability of the SNE-FM to enhance drug permeation in vitro and in vivo.

(1)

t = duration of the permeation, Mt/M∞ = accumulate percentage of permeation at t, k = rate constant of permeation, n = diffusion index of the concerned DDS. It was inferred that among L90@S10@D10, L67@S33@D10, and L50@S50@ G


Research Article


Figure 6. Accumulative permeation−time curve. (a) plotted by permeation percentage, (b) plotted by permeation amount.

Drug permeation enhancement was first tested in Franz cells with ex vivo rat abdomen skin as described in the Experimental Section. As indicated by the data shown in Figure 6 and Table 1, IBU delivered with L67@S33@D10 showed a significant

in the following 4 h, the permeation rate of L67@S33@D10 showed a significant increase compared to that of the drug solution, indicating the permeation enhancing effect of the nanoemulsion produced by the material. The time lag of L67@ S33@D10 was reduced by approximately 3 times, correlated by the increase in diffusion coefficient. These ex vivo results suggested an enhancement in drug permeation in rat abdomen skin with a notable rate which is in accordance with previous reports.30 The drug permeation enhancement was illustrated visibly in an artificial connective tissue based on a photocrosslinked gelatin hydrogel (GelMA). The hydrogel bulk consisted of a three-dimensional gelatin fibrous network and the water confined within the network, subtly simulating the extracellular matrix of the connective tissue.42 Taking such a simulating tissue as the object of drug delivery, the resulting drug distribution in the bulk could act as a prediction of the drug permeation outcome in vivo. Since the drug permeation into the tissue was directly relative to the drug penetration into the cells, this result was also a representative of the therapeutic effect. For the sake of observation, a hydrophobic drug with

Table 1. Flux, Permeation Coefficient, Time Lag, and Diffusion Coefficient of SNE-FMs Compared to PBS Solution of IBUa

L67@S33@D10 L100@S0@D10 10% IBU/PBS

Flux

Permeation Coefficient

Time Lag

Diffusion Coefficient

0.0125 0.0069 0.0096

0.012872 0.007188 0.00768

0.2 0.64 0.57

64.5333 20.1667 22.64327

Calculated by the fitted permeation curve.

a

increase in flux, permeation coefficient compared to that of L100@S0@D10 (1.79 times) and drug solution in phosphatebuffered saline (PBS) (1.67 times). Notably, the permeation− time curve of drug solution was very close to that of L67@ S33@D10 in the first 4 h when plotted by amount. However,

Figure 7. Artificial connective tissue slices under a microscope excited with green fluorescence after absorbed drug from different SNE-FMs for different durations. (a) L100@S0@D1, (b) L67@S33@D1 at 0.5 h; (c) L100@S0@D1, (d) L67@S33@D1 at 2 h; (e) V100@S0@D1, (f) V67@ S33@D1 at 5 min, and (g) V100@S0@D1, (h) V67@S33@D1 at 2 h. Drug permeation results of (i) DOX solution in UPW and (j) DOX-laden nanoemulsion are presented as control. (k) Scheme of the permeation enhancement from the SNE-FM. H


Research Article


Figure 8. Therapeutic results of the SNE-FMs in the rheumatoid arthritis model in rat induced with Freund’s adjuvant. (a−i) Macroview of the inflammation-induced paws (within the yellow box) before and after treatment. (j) Semiquantitative arthritis index evaluation on the inflammationinduced paws. (k) Thickness of the paws measured before and after the treatment (#: >thickness at 0 h, P < 0.05; *:

Self-Nanoemulsifying Electrospun Fiber Enhancing Drug Permeation

Recommend Documents