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B: Biomaterials and Membranes

Molecular Insight into Drug-Loading Capacity of PEG-PLGA Nanoparticles for Itraconazole Natalia Wilkosz, Grzegorz #azarski, Lubomír Ková#ik, Patrycja Gargas, Maria Nowakowska, Dorota Jamróz, and Mariusz Kepczynski J. Phys. Chem. B, Just Accepted Manuscript • DOI: 10.1021/acs.jpcb.8b03742 • Publication Date (Web): 21 Jun 2018 Downloaded from http://pubs.acs.org on June 23, 2018

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The Journal of Physical Chemistry

Molecular Insight into Drug-Loading Capacity of PEG-PLGA Nanoparticles for Itraconazole Natalia Wilkosz,†,§ Grzegorz Łazarski,†,§ Lubomir Kovacik,‡ Patrycja Gargas,† Maria Nowakowska,† Dorota Jamróz,† Mariusz Kepczynski*,† †

Jagiellonian University, Faculty of Chemistry, Gronostajowa 2, 30-387 Kraków, Poland ‡

Institute of Biology and Medical Genetics, First Faculty of Medicine, Charles University Albertov 4, 128 01 Prague, Czech Republic

ABSTRACT

Nanoparticles made of amphiphilic block copolymers comprising biodegradable core-forming blocks are very attractive for the preparation of drug delivery systems with sustained release. Their therapeutic applications are, however, hindered by low values of the drug loading content (DLC). The compatibility between the drug and the core-forming block of the copolymer is considered the most important factor affecting the DLC value. However, the molecular picture of the hydrophobic drug – copolymer interaction is still not fully recognized. Herein, we examined this complex issue using a range of experimental techniques in combination with atomistic molecular dynamics (MD) simulations. We performed an analysis of the interaction between

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itraconazole, a model hydrophobic drug, and a PEG-PLGA copolymer, a biodegradable copolymer commonly used for the preparation of drug delivery systems. Our results clearly show that the limited capacity of the PEG-PLGA nanoparticles for the accumulation of hydrophobic drugs is due to the fact that the drug molecules are located only at the water-polymer interface, while the interior of the PLGA core remains empty. These findings can be useful in the rational design and development of amphiphilic copolymer-based drug delivery systems.

1. INTRODUCTION

Amphiphilic block copolymers (AmBCs, containing both hydrophobic and hydrophilic blocks) are widely used for the preparation of drug carrier systems.1 In an aqueous solution above the socalled critical aggregation concentration (CAC), AmBCs can self-assemble into various micro/nanostructures (such as globular or extended micelles or vesicles) with an inner hydrophobic core and an outer hydrophilic shell. The hydrophobic core can trap poorly watersoluble drugs, while the hydrophilic corona isolates the encapsulated drug from external media. Poly(ethylene glycol) (PEG) is one of the most commonly used hydrophilic polymers that are attached to hydrophobic blocks to enhance the biocompatibility of systems,2,3 while biodegradable polyesters, such as poly(3-caprolactone) (PCL), poly(lactic acid) (PLA), or poly(lactide-co-glycolide) (PLGA) are often used as the hydrophobic blocks.4 PLGA, a random copolymer of glycolic acid (GA) and lactic acid (LA), is a biodegradable synthetic polyester that is physically stable and easily processable.5 Therefore, this copolymer is an important polymer


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matrix used for delivery and sustained release of bioactive substances through hydrolytic degradation in the body.6 PEG-PLGA block copolymers are especially widely used in biomedicine, as both PEG and PLGA segments are FDA approved for clinical use. Hence, they have been tested as biomaterials for the delivery and sustained release of proteins and hydrophilic and hydrophobic drugs.1 However, the values of drug loading contents (DLC, defined as the ratio of weight of drug enclosed in the polymer phase to the polymer weight) reported in the literature are rather low, reaching only a few percent. Such values of DLC are not satisfactory for therapeutic applications and clearly demonstrate the need for improvement. The DLC values depend on both the drug solubility in water and the architecture of the PEG-PLGA copolymer used to prepare polymeric nanoparticles (NPs). For example, the loading contents of cisplatin, a hydrophilic drug, in nanoparticles formed from a series of PEG-PLGA copolymers with various ratio of hydrophilic and hydrophobic blocks were shown to be below 1% w/w.7 For doxorubicin, another hydrophilic drug, the DLC of micelles from the PEG(2000)–PLGA(23000) copolymer was ~0.51% w/w.8 Liu et al. reported that ca. 3.47% of doxorubicin can be encapsulated into vesicles obtained from a PEG-PLGA copolymer with blocks of 5000:5000 molecular weight.9 In turn, paclitaxel, a poorly water-soluble anticancer drug, was encapsulated in PEG-PLGA nanoparticles with 0.7% or 2.5% w/w loading contents, probably depending on the length of the hydrophobic block of the copolymer used to prepare nanoparticles.10,11 In this case, the extension of the PLGA segment resulted in an increase in DLC. Recently, Xu et al. reported that loading capacity of PEG-PLGA NPs for paclitaxel can be improved by appropriate blending with PLGA.12 It was shown that DLC of PEG(5000)–PLGA(45000) nanoparticles was 2.2%, while those from the PLGA(5600)/PEG(5000)–PLGA(20000) (1:1 w/w) blend were characterized by the DLC value


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of 6.9%. Hydrophobic drugs show slightly higher affinity for PEG-PLGA nanoparticles. Song et al. reported that curcumin, a hydrophobic drug, can be encapsulated in PLGA-PEG-PLGA triblock copolymer micelles with DLC of ca. 6.4% w/w.13 Factors that affect the loading capacity of AmBC structures include the chemical nature of solute, the nature and length of the core-forming block, the total copolymer molecular weight, and to a lesser extent, the nature and block length of the corona.14 However, the most important factor is the compatibility between the drug and the core-forming block of the copolymer. Therefore, understanding the polymer-drug interaction is critically important for preparation of effective polymeric systems for drug delivery. Computer modeling is excellent for shedding light into the structural and dynamical properties of copolymer systems at atomistic or molecular levels of detail.15 Because AmBCs are too large to simulate atomistically, many researchers use coarse-grained (CG) simulations, in which several atoms are grouped into ‘beads’ that represent molecular fragments. For example, CG simulations were used to study the incorporation of paclitaxel into PEG-PLA16 and PEG-PCL nanostructures.17 Both studies have shown that the drug is evenly distributed in the hydrophobic core. However, depending on the level of coarsegraining, the beads may retain a certain amount of chemical specificity of the atoms, or they can represent only very generic properties,15 which can be important in studies of the polymer-drug interaction. To date, no detailed studies have been published on the interaction between PEG-PLGA and drug molecules. Here, we study this complex issue using a number of experimental techniques (including Langmuir film balance, and fluorescence experiments) in combination with computational atomistic molecular dynamics (MD) simulations. This multidisciplinary approach is largely the method of choice to explore nanoscale phenomena in biomolecular and


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macromolecular systems. This approach has been previously proven as very successful in generating new knowledge to develop liposome-based drug delivery systems.18,19 In this study, a PEG-PLGA copolymer was used to prepare a drug-carrier system for itraconazole (ITZ). ITZ was applied as a model hydrophobic drug. ITZ is an orally administered triazole antifungal agent that has been approved to treat serious mycotic infections in both normal and immunocompromised patients.20 However, its in vivo bioavailability is strongly limited by its low water-solubility.21 Therefore, various polymeric formulation have been proposed to improve ITZ bioavailability.22,23 We prepared the PEG-PLGA nanoparticles by a solvent-switching technique. The size of the nanoparticles was examined using dynamic light scattering (DLS) and cryo-transmission electron microscopy (cryo-TEM) was used to visualize their morphology. Next, we focused on interactions between ITZ and the PEG-PLGA NSs. The drug affinity for the polymer phase was evaluated using spectrofluorimetric measurements. The mutual miscibility of ITZ and PEG-PLGA was investigated by Langmuir monolayer experiments. Finally, we performed all-atom MD simulations to explain the drug-polymer interactions at the molecular level. The ITZ – PEG-PLGA systems studied in this paper can serve as a model for other hydrophobic drugs interacting with the PLGA matrix. These results can also be useful in searching for general rules governing the behavior of other small organic molecules, applied as drugs, in nanostructural polymer-based drug delivery systems.

2. MATERIALS AND METHODS 2.1. Materials. Poly(ethylene glycol) methyl ether-block-poly(lactide-co-glycolide) (PEGPLGA, Figure 1) was purchased from Sigma-Aldrich. The average Mn of the PLGA block was 11,500 Da and the lactide/glycolide molar ratio was 50/50. The PEG block had the average Mn of


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2000 Da. Itraconazole (ITZ), dimethylformamide (DMF), and tetrahydrofuran (THF, ≥99.9%) were obtained from Sigma-Aldrich. All experiments were conducted under phosphate-buffered saline (PBS tablets, Sigma-Aldrich) conditions, at pH 7.4. Water was purified using a water purification system (Simplicity UV, Millipore S.A.S., France).

Figure 1. Chemical structures of the poly(ethylene glycol) methyl ether-block-poly(lactide-coglycolide) copolymer (PEG-PLGA) and itraconazole (ITZ). Hydrophobic and hydrophilic blocks of the copolymer are marked in orange and black, respectively.

2.2. Apparatus.

Emission

spectra

were

measured

with

a SLM

Aminco

8100

spectrofluorimeter equipped with a 450 W xenon lamp and an external Julabo F25 system to control the temperature of the measuring chamber with an accuracy of 0.1 °C. The recorded spectra were corrected using the function provided by the manufacturer. The spectra were measured using a quartz cuvette with an optical path of 1 cm. Dynamic light scattering (DLS) measurements were performed using a Malvern Nano ZS at an angle light scattering of 173° as previously described.24 The results were analyzed using a software provided by Malvern.


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For cryo-TEM observations, samples were prepared by plunge-freezing.25 3 µL of the sample were applied on a TEM grid covered with a perforated film (lacey carbon grids #LC-200 Cu, Electron Microscopy Sciences, Hatfield, PA, USA), excess sample was removed with tissue paper (Whatman) and then the grid was immersed into liquid ethane held at –183° C. Cryo-TEM visualization was carried out using a Tecnai G2 Sphera 20 electron microscope (Thermo Fisher Scientific Inc., USA) equipped with a cryogenic Gatan 626 holder. Images were recorded at the accelerating voltage of 120 kV using a Gatan UltraScan 1000 CCD camera and a low-doseelectron imaging not exceeding 15 electrons per Å2. The layer thickness of the sample was between 100 and 300 nm. The applied magnifications resulted in pixel sizes ranging from 1 to 0.4 nm, and the typical value of the underfocus ranged from 1.5 to 2.5 µm. The surface pressure – molecular area (π – A) isotherms were recorded on a KSV NIMA 5000 Langmuir trough (KSV Instruments Ltd., Helsinki, Finland). Stock solutions of PEG-PLGA and ITZ were prepared in chloroform (cPEG-PLGA = cITZ = 0.883 mg/mL). The deionized water was used as a subphase and its temperature was kept at 25 ± 0.1 °C using a thermostat. The monolayers were left for 10 min and then compressed at a barrier speed of 10 mm/min in all experiments. The isotherms for each of the investigated monolayers were recorded at least twice, each time from a freshly prepared mixture. The reproducibility of the results was at least ±0.1 mN/m and ±0.2 nm2/molecule. 2.3. Preparation of polymer nanoparticles (NPs). PEG-PLGA (12.5 mg) was dissolved in 1 mL of THF by vortex mixing. 200 µl of deionized water was dropwise added to the PEG-PLGA solution. The prepared sample was stirred for 2 hours. After this time, 5 mL of water was added and the mixture was dialyzed against water to remove the organic solvent. NPs containing ITZ


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were prepared using the same procedure. The appropriate amount of ITZ was added to the PEGPLGA solution in THF. 2.4. Atomistic MD Simulations. System Preparation. The simulated systems are summarized in Table 1. The PLGA block was modeled by an oligomer chain consisting of 20 units. The molar ratio of glycolide to lactide in the chain was 1:1 with both L and D lactide stereoisomers at an equal ratio. Four oligomers with a random sequence of the units were generated. The oligomers were parameterized with the all-atom OPLS-AA force field,26 using common parameters designed for esters (Figure S2 and Table S1). The ITZ molecule was parameterized using the same force field, as described previously.18 The TIP3P model was used for water.27 All systems were simulated in the NpT ensemble at 310 K maintained by the velocity-scaling algorithm. The isotropic pressure of 1 bar was controlled with the Parrinello-Rahman barostat.28 Electrostatic interactions were calculated using the PME method with a cutoff radius of 1 nm for the short range.29 Van der Waals interactions were calculated with a cutoff distance of 1 nm. The simulations were performed using GROMACS version 5.0.7.30 Pure PLGA. A box with nine PLGA oligomers arranged in a square grid was generated (Figure S3). The oligomers were solvated by ca. 20,000 water molecules (PLGA system) and the whole box was energy minimized. This step was followed by a 60-ns production simulation run with the 2-fs time step. The simulation resulted in the formation of an oligomer aggregate (a polymer coil). PLGA with ITZ. We performed simulations of three systems that differ in ITZ content (12 and 24% w/w). In the case of a PLGA-ITZ1 system, the initial geometry was based on the final configuration obtained in the simulation of the PLGA system. The PLGA aggregate was placed centrally in a simulation box of a size sufficiently large to accommodate four ITZ molecules at


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a considerable distance from the PLGA aggregate. The IZT molecules were located at the apices of a regular tetrahedron around the PLGA coil (Figure S6). This arrangement was chosen to maximize the distance between the ITZ molecules, and thus to minimize the risk of their dimerization, which might hinder their entry into the PLGA aggregates. The system was rehydrated with ca. 60,000 water molecules, energy minimized and subjected to a 60-ns production run with the 2-fs time steps. A similar system, differing in the orientation of the ITZ molecules, which were rotated by 180 degrees around their principle axis, was generated and subjected to the same simulation protocol. The system corresponding to the higher ITZ content (PLGA-ITZ2 system) was generated based on the final configuration of the PLGA-ITZ1 system, to which four additional itraconazole molecules were inserted at the same initial positions (Figure S7). For the PLGA-ITZ3, the initial geometry was the same as for the PLGA system, but four ITZ molecules were placed between PLGA oligomers (Figure S8).

Table 1. Summary of simulated systems. For each system, the table indicates the number of molecules in the given system and time of the simulation.

a

System

PLGA

ITZ

Content of ITZ Water (% w/w)a

Simulation length (ns)

PLGA

9

-

0

20,000

2 × 60

PLGA-ITZ1

9

4

12

60,000

2 × 60

PLGA-ITZ2

9

8

24

60,000

2 × 120

PLGA-ITZ3

9

4

12

60,000

2 × 120

The ITZ content is a weight fraction of the drug relative to the oligomer weight.

3. RESULTS


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3.1. Size and Morphology of PEG-PLGA Nanoparticles. NPs of the PEG-PLGA copolymer were fabricated by the solvent-switching technique previously described by Soo and Eisenberg.31 The as-prepared NPs were examined using the DLS technique. The mean hydrodynamic diameter (dz) of the objects present in the PEG-PLGA dispersion was about 142 nm and the polydispersity index (PDI) was ~0.24. The distribution profile of the hydrodynamic diameters is shown in Figure S1. These results indicate that the nanometric structures can be obtained from the copolymer, but the population of objects was polydisperse. For further analysis of the PEG-PLGA dispersion, we used direct observation with cryo-TEM microscopy. Figure 2 shows typical cryo-TEM micrographs of the PEG-PLGA dispersion. The images reveal the presence of two populations of objects: smaller solid particles, which can be attributed to polymeric micelles formed in the system, and hollow structures of larger sizes. The presence of the latter indicates that the copolymer can also spontaneously self-organize into polymersomes, spherical vesicles with a distinct polymer membrane surrounding the aqueous phase. The cryo-TEM microscopy observations allowed us to distinguish the empty and solid objects and to determine size statistics for both types of nanostructures. Distribution profiles are presented in Figure 2B. The size distribution of the micelles is narrow and the average diameter was 42 ± 15 nm, while the average diameter of the polymersomes was 123.0 ± 32.0 nm. In addition, on the basis of microscopic observations at high magnification, the average thickness of the polymer membrane was estimated to be 30.0 ± 6.7 nm. The microscopic observations are consistent with the DLS results. It should be noted that the DLS technique provides data on the mean hydrodynamic diameter that is heavily weighted toward the largest structures present in the system.32,33 Thus, DLS shows systematically larger mean sizes of the objects studied.


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It is known that the molecular size, composition, architecture, and concentration of AmBCs affect the size and morphology of assemblies formed in an aqueous solution.34,35 Self-assembly of PEG-based block copolymers in a diluted aqueous solution has been extensively studied. It was found that changes in the balance of hydrophilic/hydrophobic segments (referred to as PEG weight fraction, wEO) have a significant impact on the morphology of the objects observed in polymer dispersions; spherical and worm-like micelles, vesicles, and solid-like particle can be formed depending on the wEO value. However, the ranges of wEO, in which the specific NPs are stable are also depended on the type of hydrophobic block. For example, diblock copolymers containing poly-(1,2-butadiene) (PEG-PBD) or poly(ethyl ethylene) (PEG-PEE) can form polymersomes when the hydrophilic fraction is in the range of 0.25-0.45.35,36,37 When the PEG content increases, the polymer produces cylindrical and spherical micelles. In contrast, copolymers based on biodegradable PCL (PEG-PCL) can self-assembly into polymersomes when wEO is ~0.15.38 The branched copolymers composed of three PEG chains linked to a PLA chain ((PEG)3-PLA) were shown to form vesicles at the hydrophilic fractions in the range of 0.1 – 0.3.2 The wEO value of the PEG-PLGA polymer used in the current studies is about 0.15 and we observed the coexistence of small micelle-like structures and polymersomes in a diluted dispersion of this polymer. Studies on the nature of structures formed by the PEG-PLGA diblock copolymers are rather scarce in the literature. They indicate, however, that beside the hydrophilic/hydrophobic block ratio, also the preparation method affects the size and morphology of NPs. For example, it was shown that using the dialysis method micelles with an average size of 26 nm can be obtained from the copolymer with the hydrophilic fraction of 0.155, very similar to that characterizing the polymer used in our study.13 Avgoustakis et al. investigated a series of PEG-PLGA copolymers with wEO varying between 0.068 – 0.42.7 Only


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solid-like nanoparticles of around 150 nm in diameter were prepared from these polymer by the double emulsion method. In turn, the formation of polymersomes was demonstrated using cryoTEM for the PEG-PLGA copolymer with wEO equal to 0.5.9 These results clearly show that the determination of the relationship between the architecture of PEG-PLGA copolymers used to prepare the NPs and type of particle formed requires more systematic study.

Figure 2. (A) Cryo-TEM micrographs and (B) the diameter profiles (calculated from the cryoTEM data) of the PEG-PLGA nanostructures (cPEG-PLGA = 2.5 mg/mL). Scale bars correspond to 100 nm.

4.2. Interaction of ITZ with the PEG-PLGA NPs. Fluorescence Experiments Indicate Good Affinity of ITZ for the PEG-PLGA NPs. The partitioning of ITZ between the PEG-PLGA NPs and the bulk aqueous phase was investigated using the fluorescence spectroscopic titration


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technique.39 The affinity of the drug to enter the polymer phase can be quantified using a socalled binding constant, defined as follows:

Kb =

cP

(1)

cw cPEG -PLGA

where cP and cw are ITZ concentrations in the polymer and aqueous phase, respectively; cPEGPLGA

is the concentration of PEG-PLGA in the system. The partitioning of ITZ to NPs was

studied by recording steady-state emission spectra of solutions containing a constant drug concentration and PEG-PLGA concentrations varying from 0 to 0.33 mg/mL. Figure 3A presents a typical set of spectra obtained for ITZ in the presence of PEG-PLGA at various concentrations. An amplification of the fluorescence intensity was observed with increasing concentration of the copolymer. It was accompanied by a blue shift of the fluorescence band (the fluorescence maximum shifted from 372 nm to 366 nm). The binding constant was estimated based on these spectral changes, which were attributed to the transfer of ITZ molecules from the aqueous phase to the polymer phase. The effective binding constant, Kb, was determined by fitting the experimental data to the following formula (see the Supporting Information):40

F=

Finit + Fcomp K b cPEG-PLGA

(2)

1 + K b cPEG-PLGA

where Finit, F and Fcomp are the fluorescence intensity of the drug measured in the absence of the copolymer, in the presence of the copolymer at a concentration of cPEG-PLGA and the asymptotic value of the intensity at complete binding, respectively. To obtain the Kb, F versus cPEG-PLGA data were plotted and fitted to equation 2 by a nonlinear regression routine. Figure 3B shows the dependence of ITZ fluorescence intensity at 366 nm on cPEG-PLGA and the line fitted to the plot. The binding constant of ITZ to NPs was found to be Kb = 30.9 ± 5.5 mL/mg. This value shows


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that at cPEG-PLGA = 32 µg/mL 50% molecules of ITZ are embedded in the PEG-PLGA NPs. One can easily calculate that, at the polymer concentration of 1 mg/mL used in our studies, the concentration ratio of ITZ encapsulated in the NPs and dissolved in the aqueous phase was as high as 31. This result confirms that ITZ has relatively good affinity for the PEG-PLGA NPs.

Figure 3. (A) Emission spectra (λex = 260 nm) of ITZ (cITZ = 1.4 µM) in PBS buffer (pH 7.4) and in PEG-PLGA dispersions with increasing polymer concentrations. (B) The ITZ fluorescence intensity at 366 nm versus the PEG-PLGA concentration, along with the line fitted to equation 2.


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DLS Measurements Demonstrate Limitations in Loading of NPs with ITZ. DLS measurements were carried out to determine the amount of ITZ that can be incorporated into the PEG-PLGA NPs without causing changes in their morphology. A series of samples with the ITZ content up to 10.0% w/w and a constant copolymer concentration (1 mg/mL) were prepared and hydrodynamic diameters were measured. Incorporation of small amounts of the drug (up to 5.0% w/w) resulted in an increase in the size of the ITZ-loaded NPs. Figure S1 shows the distribution profile of the hydrodynamic diameters of the PEG-PLGA nanoparticles containing 5.0% w/w of the drug. The mean hydrodynamic diameter of the objects increased to about 152 nm, while PDI increased to 0.35 ± 0.02. However, the incorporation of more ITZ led to the appearance of precipitate in the dispersion after dialysis and to a drastic increase in the PDI value. Therefore, we were not able to determine the dz values. Monolayer Experiments Show the Partial Miscibility of ITZ with PEG-PLGA. The mutual miscibility of ITZ and PEG-PLGA is of fundamental importance for effective embedding of the drug into the NPs. We used Langmuir balance measurements and Brewster angle microscopy (BAM) to study their miscibility. π−A isotherms measured during the compression of one component (PEG-PLGA or ITZ) films, and their mixtures of various compositions are shown in Figure 4. In the case of the pure copolymer monolayer, the surface pressure begin to increase at a molecular area of approximately 40 nm2. Then, π increases gradually during the compression up to a value of ∼11 mN/m, at which it levels off forming a plateau before a sharp increase to the A value of ∼2.5 nm2. The presence of this plateau can be attributed to the desorption of PEG blocks from the air−water interface to the aqueous phase. The isotherm shape for the PEG-PLGA monolayer is very similar to that previously observed for various PEGylated lipid systems.18,41 For these systems, the plateau corresponding to the removal of PEG chains from the interface


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was observed at the similar surface pressure (∼13 mN/m). The admixture of ITZ to the PEGPLGA monolayer essentially shifts the isotherms toward smaller molecular areas. The drug has a little effect on the polymer desorption to water, since the inflection in the isotherms is almost at the same π for all ITZ mole fractions used in the experiment. As can be expected, the plateau length becomes shorter with decreasing polymer content in the films.

Figure 4. Selected surface pressure (π)−area (A) isotherms recorded for the PEG-PLGA/ITZ monolayers on the water subphase at 25°C. XITZ stands for the mole fraction of ITZ.

BAM images taken during compression of the mixed PEG-PLGA/ITZ films are presented in Figure 5. In the case of pure copolymer and systems with XITZ ≤ 0.7, the monolayers are homogeneous up to the collapse surface pressure. In contrast, for the ITZ film, bright oval domains start to appear at π ∼ 4.5 mN/m. These domains grow and begin to merge during the film compression. Very high brightness of these domains may indicate that they are multilayer structures.18 For the monolayer with XITZ = 0.94, small bright spots can be observed at π = 7.5 mN/m and the further π increase causes the bright structures to become larger.


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The results from the π−A isotherms and BAM microscopy show that the PEG-PLGA monolayers containing up to 70 mol% of ITZ are completely homogeneous indicating that ITZ mixes with the copolymer in that concentration range. Due to the large difference in molecular weights between PEG-PLGA and ITZ, the value of XITZ = 0.7 corresponds only to 12.2% w/w. Increasing the ITZ concentration in the PEG-PLGA monolayer causes the appearance of bright structures in the BAM images, and they persist up to the pressure of collapse. The features are very similar to those observed for the one-component ITZ film. This suggests that at higher drug concentrations, the interactions between PEG-PLGA and ITZ are unfavorable and lead to phase separation with ITZ expelled from the copolymer film.


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Figure 5. Brewster angle microscopy (BAM) pictures of the PEG-PLGA/ITZ monolayers taken at different stages of film compression at 25 °C.


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4.3. Molecular View of the ITZ-PLGA Interactions. To obtain a molecular insight into the interaction between ITZ and the NP core created by the PLGA blocks, we carried out atomistic MD simulations of four model systems (Table 1). First, we simulated systems containing nine PLGA oligomers immersed in the water phase (PLGA system). Snapshots taken at the beginning and at the end of the simulations showing the conformation of the oligomers are presented in Figure S3. As can be seen, all the oligomers present in the box aggregated to form a tightly packed amorphous polymer coil. The aggregation process was completed within the first 1.5 ns of the calculations. The polymer coil did not include any water molecules, as can be demonstrated by visual inspection of the coil cross-sections (Figure 6) and calculating the cumulative number of water molecules as a function of a distance from the center of mass of the PLGA coil (Figure S4). This proves that the PLGA core is strictly hydrophobic since the first water molecules are encountered at ca. 1.3 nm from the core center, which is a distance comparable to the approximate radius of the oligomer aggregate (ca. 1.5 nm, see the inset of Figure S4). Our simulation results are consistent with experimentally determined hydration numbers of some esters to be ca. 0, suggesting that there is no hydrogen bond formation between the ester group and water molecules.42 To further compare the outcome of our simulations with experimental results, we calculated the mass density of the PLGA coil. As shown in Figure S5, the polymer aggregate is rather homogeneous to a distance of about 1.2 nm from its center of mass and the average density the hydrophobic core is about 1.36 g/cm3. This value is within the range of experimentally determined densities (1.26–1.58 g/cm3) for PLGA nanoparticles with different lactide/glycolide molar ratios.43,44 Therefore, it can be assumed that system PLGA is a good model of the hydrophobic core of the PEG-PLGA NPs.


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Figure 6. Cross-sections of the PLGA coil in the yz (left) and xz (right) planes. All the water molecules found within 0.4 nm of any atom of the coil are shown as cyan surfaces. The oligomer chain is represented in yellow, with oxygen atoms marked in red and hydrogen atoms in white.

To study the partitioning of ITZ between the aqueous phase and the PLGA core, three systems were simulated. In two systems, ITZ was added to the preexisting PLGA coil at various weight ratios, closely mimicing the experimental setup, where the drug was introduced to NPs. In the case of the 12.2% w/w drug content, four ITZ molecules were incorporated into the PLGA system at the most distant corners of the simulation boxes to avoid ITZ aggregation (the PLGAITZ1 system, Figure S6). Figure 7A depicts how the distance of the ITZ molecules from the center of mass of the polymer coil changed during simulation. As can be seen, all the ITZ molecules migrated to the vicinity of the polymeric coil within approximately 25 ns. For the rest of the duration of the simulation, ITZs resided near the water-PLGA interface and did not display the tendency to translocate to the core of the polymer coil. A selected image taken at the end of the simulations, illustrating the preferred position and orientation of the ITZ molecules in the PLGA aggregate, is shown in Figure 8A. It clearly demonstrates that ITZs align with their long


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axis almost parallel to the polymer surface and they are not evenly distributed on the surface of the coil. To understand the effect of the ITZ content on its behavior in the PLGA phase, we added four ITZ molecules to the equilibrated PLGA-ITZ1 system (Figure S7), which corresponds to the drug content of 24% w/w. Figure 7B shows the time profile of the distance between the centers of mass of the ITZ molecules and that of the polymer coil. It demonstrates that after ca. 100 ns all the additional drug molecules migrated to the water-polymer interface and adhered to the PLGA coil. The final configuration of the PLGA-ITZ2 system is presented in Figure 8B. The uneven distribution of the drug molecules is clearly visible. The ITZ molecules gathered together to form an aggregate that is attached to the surface of the PLGA coil. This finding is consistent with the monolayer experiments showing that at higher drug concentrations ITZ was expelled from the PEG-PLGA copolymer film and the phase separation occurred.

Figure 7. The time profiles of the distance of the centers of mass of the ITZ molecules and the PLGA coils in the PLGA-ITZ1 (A) and PLGA-ITZ2 (B) systems.


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In the initial configuration of the PLGA-ITZ3 system, four ITZ molecules were inserted between nine PLGA oligomers arranged analogously to the PLGA system (Figure S8). This corresponds to inclusion of the drug at the stage of NP formation during dialysis. During the simulation, the PLGA chains aggregated within 60 ns, thus forming a polymer coil. Concomitantly, the drug molecules were expelled from the coil and deposited at its surface, as shown in Figure 8C. Concluding, in all systems itraconazole molecules showed a clear tendency to accumulate only on the surface of the PLGA coil. The initially free ITZ molecules migrate relatively quickly to the coil surface and, once anchored in this position, they do not tend to move to the core of the polymer aggregate. For that reason, the volume of PEG-PLGA nanoparticles available for the drug is very limited, as a result their drug loading capacity is low.

Figure 8. The final configuration of the PLGA-ITZ1 (A), PLGA-ITZ2 (B), and PLGA-ITZ3 (C) systems.


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4. DISCUSSION In this paper, we address the hydrophobic drug – PEG-PLGA matrix interaction in terms of drug loading capacity of the copolymer NPs. It is known that the nature of the drug, including polarity, hydrophobicity, charge and degree of ionization, greatly affects its incorporation to polymeric matrices; however, the compatibility between the drug and the core-forming segment makes incorporation most effective.14 One of the parameters proposed to predict the polymerdrug compatibility is the Flory–Huggins interaction parameter. This interaction parameter (χdp) between the drug (d) and the hydrophobic block (p) can be described by the following equation:23

߯ୢ୮ = ൫δୢ − δ୮ ൯

ଶ ௏ౚ ோ்

= ሺ∆δሻଶ

௏ౚ

(3)

ோ்

where δd and δp are the Hildebrand solubility parameters of the drug and the hydrophobic block, and Vd is the molar volume of the drug. Differences in the solubility parameter for drug and polymer (∆δ) can be treated as a measure of compatibility (miscibility) of both ingredients. It has been proposed that ∆δ values suggesting the drug – polymer miscibility should be less than 2 MPa0.5.45,46 The solubility parameters of PLGA with the lactide/glycolide molar ratio of 50/50 have been determined by various methods and ranged from 19.9 to 23.1 MPa0.5.47,48 The average value calculated on the basis of the reported values is 21.2 ± 1.5 MPa0.5. For ITZ, the δd values were reported in a range of 20.81 – 22.6 MPa0.5.22,23,49 Based on these values, one can conclude that ∆δ for the ITZ-PLGA pair is below the threshold value of 2 MPa0.5, suggesting good miscibility of these two compounds.


23


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Our spectrofluorimetric experiments confirm the high affinity of ITZ (at the low concentration) to the PLGA phase. This is an expected result, since ITZ as a weak basic drug with pKas ≤ 4.0,50 is unionized under our experimental conditions and its neutral form is characterized by very poor water-solubility21 and high lipophilicity (log P, a lipophilicity parameter, is 6.2). Therefore, ITZ molecules escape from the aqueous phase and partition into the polymer phase, as indicated by the Kb value. On the other hand, the monolayer measurements clearly show that ITZ is miscible in the PEGPLGA copolymer only up to DLC of about 12% w/w. At higher drug concentrations the phase separation takes place and ITZ is expelled from the monolayer forming multilayer stacks, as was shown using the BAM observations. To further analyze the interactions between ITZ and PEGPLGA we calculated the values of excess areas of mixing (AExc) based on the π−A isotherms according to the equations:51

A୉୶ୡ = A − A୧ୢ = A − ሺ୍ܺ୘୞ A୍୘୞ − ܺ୔୉ୋି୔୐ୋ୅ A୔୉ୋି୔୐ୋ୅ ሻ,

(4)

were A is the mean area per molecule in the mixed monolayer derived from the isotherms at a given surface pressure, where Aid is the molecular area corresponding to ideal mixing. Xi and Ai are a mole fraction in the mixed films and the mean areas per molecule of the ith components in their pure films at a given π, respectively. Thus, AExc is a measure of deviation of the mean molecular area determined from the isotherm of the mixed film from the area per molecule calculated assuming an ideal mixing of two components. The calculations were performed at π = 5 mN/m, which is the highest surface pressure at which ITZ still forms the monolayer. The AExc values were plotted as a function of the film composition (Figure 9) to evaluate the magnitude of


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the interaction between the molecules in the mixed systems. As can be seen, the values of AExc were positive for all the mixed PEG-PLGA/ITZ compositions, which indicates that the intermolecular interactions between the drug and the copolymer are less favorable compared to those in the one-component systems.

Figure 9. Excess areas of mixing (AExc) as a function of the monolayer composition calculated at the surface pressures of 5 mN/m.

Our simulation results are consistent with the experimental findings. They demonstrate that all the drug molecules initially placed in water readily translocate to the polymer coil. However, due to the unfavorable interactions between the copolymer and the drug, ITZ does not penetrate deeply into the polymer phase, but remains adhered to the surface. Such location of ITZ in the PLGA aggregate appears to be justified considering the chemical structures of the copolymer and the drug and its interactions with water. First, all polar groups of both ITZ and PLGA are hydrogen acceptors (see Figure 1), therefore there is no possibility of direct hydrogen bonding between them. Second, the ITZ molecule has several electronegative atoms in its chemical structure (8 N atoms, 4 O atoms, and 2 Cl atoms). These atoms are almost evenly distributed


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along the molecular skeleton and in our simulations they bear partial charges ranging from −0.066 e to −0.61 e. Additionally, we have shown previously that the ITZ molecule can form about 2.8 hydrogen bonds with water.3 All this makes the ITZ molecules expelled from the interior of the hydrophobic core of the PLGA coil to the optimal position at the surface. This location of ITZ reduces the contact between ITZ hydrophobic groups (phenyls, 1-methylpropyl) and water, on the one hand, and allows hydrogen bonding between its electronegative atoms and water molecules on the other hand, what results in additional stabilization of the system. In our simulations, we considered only the hydrophobic core of the PEG-PLGA NPs. It is, however, surrounded by the hydrophilic shell. Important issues related to the presence of the PEG corona are (i) whether PEG penetrates into the PLGA phase, changing its properties, and (ii) whether the PEG chains affect the accumulation of ITZ in the NPs. The partitioning of PEG blocks into a hydrophobic core was previously studied using the CG simulations. In the case of PEG-PLA, it was demonstrated that the hydrophilic blocks did not enter the hydrophobic parts of the nanoparticles.16 In turn, Loverde et al. observed that PEG penetrates into the interface of the PEG-PCL micelles.17 After the incorporation of paclitaxel into the micelles, PEG enters deeper into the PCL core and its density slightly decreases. The effect of the PEG corona on the accumulation of hydrophobic compounds has been studied for PEGylated liposomes (liposomes stabilized by PEG chains covalently attached to lipids).3,18 ITZ and porphyrins have been shown to partition into both the lipid membrane core and the PEG corona. PEG chains wrap around the hydrophobic molecules, thus the PEG corona provides an additional volume for accumulation of hydrophobic compounds. All these considerations indicate that the PEG chains surrounding carrier particles have a beneficial effect on their drug loading capacity.


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6. CONCLUSIONS

In this study, we performed an analysis of the interaction between ITZ, a water insoluble drug, and the polymer matrix composed of the PEG-PLGA copolymer, considered as a possible drug carrier. A synergistic combination of experimental methods and atomistic MD simulations was used to obtain detailed information on the drug accumulation in the copolymer nanostructures. Our experimental results suggest that ITZ has a relatively good affinity for the PEG-PLGA NPs likely due to its hydrophobic nature. However, its miscibility with the core-forming block is limited due to the unfavorable interactions between these two components. In line with experimental findings, the MD simulations indicate that ITZ molecules tend to locate at the water-PLGA interface rather than in the highly hydrophobic PLGA core of NPs, reflecting the role of charge distribution in the ITZ molecule and the lack of direct hydrogen bonding between these two components. As a result of the peripheral location of the drug molecules, the volume of NPs available for their accumulation is limited only to surface layers, lowering the drug loading capacity of the copolymer structures. Therefore, we believe that the value of drug loading content for ITZ is strongly related to the surface of the AmBC structures. In summary, due to their biodegradability and biocompatibility, the PEG-PLGA nanostructures are very attractive polymeric matrices for drug delivery purposes. The problem with their relatively low drug loading capacity associated with the fact that drugs are accumulated only at the water-nanoparticle interface could be diminished by creating polymeric structures with a developed surface area, e.g. polymersomes or small micelles instead of larger polymer particles. Another possibility is the incorporation of additives that would increase a free volume of the PLGA core (an unoccupied volume enclosed in the hydrophobic core), thus reducing the high density of the PEG-PLGA NPs and improving their DLC, as demonstrated for liposomal drug


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carriers.39 Our results also show that further studies using comprehensive experimental and computer simulation methods for other drug-polymeric matrix systems could be necessary to obtain a mechanistic insight into the rational design of drug delivery systems based on AmBC self-assembles.

ASSOCIATED CONTENT Supporting Information. The Supporting Information is available free of charge on the ACS Publications website. The following files are available free of charge. Additional figures for DLS measurements and MD simulations, table with OPLS-AA atom types and partial charges used in the simulations (PDF)

AUTHOR INFORMATION Corresponding Author *E-mail: [email protected]. Tel.: +48 12 686 2532. Fax: +48 12 686 2750 (M.K.). ORCID Mariusz Kepczynski: 0000-0002-7304-6881 Author Contributions


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The manuscript was written through contributions of all authors. §(N.W. and G.Ł.) These authors contributed equally. Notes The authors declare no competing financial interest. ACKNOWLEDGMENT The project was financed by the National Science Centre,

Poland (grant nr.

2016/21/B/ST5/00250). The MD simulations were carried out using the PLGRID infrastructure (the Prometheus cluster). N.W. acknowledges the National Science Centre, Poland (NCN) for financial support in the form of ETIUDA doctoral scholarship on the basis of the decision number DEC-2017/24/T/ST5/00383. L. K. acknowledges the support of the P302/12/G157 grant of the Czech Science Foundation.

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