Synthesis and micellization of a pH-sensitive diblock copolymer for drug delivery

International Journal of Pharmaceutics 455 (2013) 5–13

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International Journal of Pharmaceutics journal homepage: www.elsevier.com/locate/ijpharm

Pharmaceutical nanotechnology

Synthesis and micellization of a pH-sensitive diblock copolymer for drug delivery Konstantinos P. Koutroumanis, Richard G. Holdich, Stella Georgiadou ∗ Department of Chemical Engineering, Loughborough University, Loughborough, Leicestershire LE11 3TU, UK

a r t i c l e

i n f o

Article history: Received 22 April 2013 Received in revised form 22 June 2013 Accepted 27 June 2013 Available online 9 July 2013 Keywords: Block copolymers Micelles pH-sensitive drug delivery Poly(ethylene glycol) Hydrazone Polycaprolactone

a b s t r a c t A diblock copolymer constituting of a poly(ethylene glycol) (PEG) and a polycaprolactone (PCL) segment, linked with a pH-sensitive hydrazone bond (Hyd), was synthesized. Micelles formed from this copolymer, offer the advantage of encapsulating hydrophobic drugs without the need for conjugation sites. All synthesized polymers were characterized using gel permeation chromatography, infrared and proton nuclear spectroscopies. PEG-Hyd-PCL micelles were prepared using the solvent-displacement method and ␣-tocopherol was used as a model drug due to its high hydrophobicity. The micelle size and drug loading efficiency were studied with regards to the hydrophilic ratio, f, molecular weight, and the polymer/drug ratio. Dynamic light scattering and transmission electron microscopy showed that the PEG-Hyd-PCL micelles had sizes ranging from 50 to 200 nm. Aqueous micellar dispersions exhibited significantly higher values of turbidity in mildy acidic pH than in neutral, indicating pH-sensitivity for the PEG-Hyd-PCL micelles. The zeta potential of the micellar solutions decreased and the molecular weight distribution became bimodal at mildly acidic pH also supporting the pH sensitive properties of the copolymer. The critical micelle concentration was calculated using fluorescence spectroscopy. © 2013 Elsevier B.V. All rights reserved.

1. Introduction Nanoparticles are one of the preferred vesicles for developing pharmaceutical formulations, mainly due to their size, as sizes smaller than 200 nm remain undetected by reticuloendothelial systems, RES (Grislain et al., 1983), achieving long systemic circulation and thus advantaging from the enhanced permeation and retention, EPR, effect (Maeda, 2001). Amongst numerous nanomaterials used in nanoparticulate drug delivery applications (Sahoo and Labhasetwar, 2003), self-assembled block copolymers combine both exceptional biocompatibility and responsiveness to bioenvironmental factors, such as pH (Letchford and Burt, 2007). pH is widely used as a trigger stimulus for targeted drug delivery, since its variations in the human organism are significant. However, for delivering a drug to cancer cells, pH-sensitivity must be achieved for only a very small variation in pH: from 7.4 in blood stream, to 6.8 in cancer cells and 5–6 in intracellular space. This difficult task is today met by micellar systems that rely mostly in two strategies: 1) destabilization of the micelles due to polarity change and subsequent release of the core physicochemically entrapped drug, or

∗ Corresponding author. Tel.: +44 1509 222521; fax: +44 1509 223923. E-mail addresses: [email protected], [email protected] (S. Georgiadou). 0378-5173/$ – see front matter © 2013 Elsevier B.V. All rights reserved. http://dx.doi.org/10.1016/j.ijpharm.2013.06.071

2) hydrolysis of an acid-sensitive bond, used to conjugate a drug to the polymer. Micelle destabilization can be achieved by using acid-sensitive groups on the hydrophobic core forming block, i.e. amino groups. Ionization at acidic pH renders the hydrophobic block hydrophilic, destabilizing the micelle, causing it to collapse, releasing the entrapped drug (Martin et al., 1996; Lee et al., 2003). Destabilization can also be the result of chain degradation. Acetal groups that were introduced by modification into the hydrophobic block of poly(ethylene glycol)-b-polyaspartic acid (Gillies and Fréchet, 2003), or poly(ethylene glycol)-b-polycarbonate (Chen et al., 2009), hydrolysed in mildly acidic pH, causing the micelles to collapse. These micelles released higher amounts of encapsulated paclitaxel or doxorubicin when pH changed from 7.4 to 5. Poly(ortho ester) side chains attached to the polymethacrylate block of a copolymer (Tang et al., 2011) have also been used and they proved to hydrolyze faster than acetal groups. Heller et al. (2002) used four types of poly(ortho esters) as the hydrophobic blocks of di- and triblock copolymers. Micelles formed were pH-sensitive and able to encapsulate paclitaxel at a high loading. Although the micelle destabilization strategy has shown promising results, it is either based on polyacrylates and vinyl polymers which have poor biocompatibility, or poly(amino acids) which biodegrade mostly enzymatically, resulting in poor controlled release in vivo (Uhrich et al., 1999). Apart from Heller’s group, now patented, work, polyesters, amongst the most important

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Fig. 1. Comparison of the new design to current acid-sensitive polymer–drug conjugates delivery systems.

biodegradable polymers, have not been utilized in systems that rely on micelle destabilization for delivering a drug, but only in systems based on the second strategy: polymer–drug conjugates. In polymer–drug conjugates, the drug is most commonly grafted onto the hydrophobic core-forming block of the copolymer. There are several links that hydrolyze in mildy acidic pH and have been used to form polymer–drug conjugates, such as carbamate, cisacotinyl and hydrazone. The cis-acotinyl and hydrazone linkages were found to hydrolyze significantly faster at acidic pH than the carbamate one (Yoo and Park, 2001; Yoo et al., 2002). The cisaconityl bond showed greater pH-sensitivity than the hydrazone one; however, its hydrolysis resulted to the release of chemically modified drug (Yoo et al., 2002). The increased pH-sensitivity of the hydrazone bond compared to the carbamate one was also confirmed by Hu et al. (2010). Conjugation requires the existence of appropriate conjugation sites on both polymer and drug. However, not all drugs have appropriate conjugation sites, which constitutes a serious issue. Modifying the drug to create conjugation sites can be a solution but adds at least one extra step to the preparation procedure (Alani et al., 2010; Bae et al., 2003; Hruby´ et al., 2005), while increasing the risk of release of chemically modified drug. Ponta and Bae (2010) used spacers between the polymer and the hydrazone-drug bond; interestingly, this affected the hydrolysis rate and the release profile of the drug. In all the above polymer–drug conjugate systems, drug release did not occur as fast as one would expect by studying the acid-liable bond hydrolysis rate. This is because the drug is freed by hydrolysis inside the core, and it has to diffuse though the micelle and to the surface to be released. In this work, we designed and synthesized a novel diblock copolymer in order to address all the above issues: the need for conjugation sites on drug and polymer, the release of chemically modified drug and the delayed drug release. The novel polymer incorporates a pH-sensitive hydrazone bond which links the two blocks. In such a system, drug simply partitions into the micellar core due to hydrophobic interactions, whereas drug release should occur due to the pH-sensitive detachment of the two blocks. Moreover, hydrolysis will result to complete breakdown of the micelles, leading to immediate drug release. Such systems could overshadow traditional micellar drug conjugates in that they would be “universal”. They provide a pH-sensitive delivery system without the

need for appropriate functional groups on each drug in order form a conjugate; any hydrophobic drug would do. Lastly, any concerns for potential release of chemically modified drug are eliminated (Fig. 1). To our knowledge, the hydrazone link has not yet been incorporated in a diblock copolymer. The only reported case is that of Kale and Torchilin (2007) who used a hydrazone bond to conjugate poly(ethylene glycol) with phosphadylethanolamine. The conjugate formed micelles which were highly sensitive to mildy acidic pH. In this work Poly(ethylene glycol), PEG, was used for the hydrophilic, shell-forming block, while ε-Polycaprolactone, ε-PCL, was used for the hydrophobic, core-forming block. PEG has exceptional physiochemical, biological properties and stealth properties, while lacking toxicity and immunogenicity (Adams et al., 2003). ε-PCL is highly biocompatible and has a very low glass transition temperature, which is essential for the formation of dynamic micelles (Munk, 1996). Moreover, ε-PCL, in contrast with other biocompatible polyesters, does not generate an acidic environment while it degrades (Sinha et al., 2004), which could potentially be harmful to acid-sensitive tissues. This work focuses on the diblock copolymer synthesis, characterization, micellization and the factors that affect the size distribution of the micelles such as molecular weight, hydrophilic fraction and drug content, and the drug encapsulation capacity. Drug release studies and profiles for various types of PEG-Hyd-PCL micelles and different conditions are on-going. 2. Experimental 2.1. Materials O-[2-(6-Oxocaproylamino)ethyl]-O -methylpolyethylene glycol (Mn = 2000 g/mol and 5000 g/mol), PEG-CHO, ε-Caprolactone monomer 99%, ε-CL, 2-hydroxyethylhydrazine 98%, 2-HEH, tin(II) 2-ethylhexanoate 95%, Sn(Oct)2 , and (±)-␣-tocopherol of ≥96% assay were purchased from Sigma–Aldrich Co Ltd, Gillingham, Dorset, UK. Toluene, extra dry grade in an AcroSeal container, was purchased from Acros Organics via Fisher Scientific Ltd, Loughborough, Leicestershire, UK. Ethanol absolute 99.8%, and all other chemicals used were of analytical grade and were purchased from Fisher Scientific Ltd, Loughborough, Leicestershire, UK.

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Fig. 2. (a) Poly(ethylene glycol) functionalization, (b) ε-caprolactone polymerization.

PEG-CHO and the hydrazone derivative dried with azeotropic distillation from toluene prior to the hydrazone link formation and polymerization reactions. Ethanol was distilled and the distillate dried over 3A molecular sieves. ε-CL monomer dried under 3A molecular sieves and distilled under reduced pressure prior to use. All other chemicals were used without any other purification.

2.5. Fourier transformed-infrared spectroscopy (FTIR)

2.2. Poly(ethylene glycol) functionalization 1 g of PEG-CHO was dissolved in 12 mL ethanol at 35 ◦ C under nitrogen atmosphere. Then, a 10% v/v 2-HEH solution in ethanol was added. 2-HEH was used in excess to ensure complete conversion of the aldehyde-terminated PEG to the hydrazone derivative (Fig. 2a). -CHO/-NHNH2 groups ratio was 1:5. The solution was left in the dark, under nitrogen atmosphere and constant stirring for 48 h. After completion of the reaction, PEG-Hyd-OH was isolated from diethyl ether. The precipitate was transferred into a −18 ◦ C freezer and washed several times with cold (−18 ◦ C) ethanol. Finally, PEG-Hyd-OH was dried under vacuum at 40 ◦ C for 24 h for complete removal of 2-HEH. 2.3. ε-Caprolactone polymerization 0.2 g of PEG-Hyd-OH and a 20% v/v ␧-CL solution in toluene were dissolved in 4 mL of refluxing toluene under nitrogen atmosphere. The mole ratio of polymer/ε-CL varied from 1:22 to 1:88 (Table 1). Polymerization (Fig. 2b) was initiated by the addition of a 10% (v/v) solution of Sn(Oct)2 in toluene (0.75 wt%) and continued at 110 ◦ C under nitrogen atmosphere and constant stirring for 18 h. The final product, PEG-Hyd-PCL, was isolated by precipitation in diethyl ether and dried under vacuum. 2.4. Gel permeation chromatography (GPC) GPC analysis was performed using an Agilent 1100 HPLC System equipped with a refractive index detector (G1362A). Analysis was performed on an Agilent PLgel MIXED-C column, 5 ␮m, Table 1 Poly(ethylene glycol) to ␧-caprolactone feed mole ratio and product coding. Feed mole ratio

Coding

PEG-Hyd-OH/ε-CL

PEG-Hyd-OH/PCL

1:22 1:44 1:66 1:88

1:0.5 1:1 1:1.5 1:2

300 mm × 7.5 mm, in series with an Agilent PLgel guard column, 5 ␮m, 50 mm × 7.5 mm. The flow rate of the mobile phase (tetrahydrofuran) was 1 mL/min and the column temperature was 30 ◦ C. Calibration was performed using polystyrene standards with a narrow molecular weight distribution (EasiVials PS-M).

PEG-Hyd-PCL-A PEG-Hyd-PCL-B PEG-Hyd-PCL-C PEG-Hyd-PCL-D

FTIR spectra were obtained using a Shimadzu FTIR-8400S spectrometer. A small amount of each material was mixed with KBr and compressed to tablets. The IR spectra of these tablets were obtained in absorbance mode and in the spectral region of 600–4000 cm−1 using a resolution of 4 cm−1 and 64 co-added scans. 2.6. Nuclear magnetic resonance spectroscopy (NMR) Polymers were solubilized in deuterated chloroform (CDCl3 ) and 1 H NMR spectra were obtained on a Bruker Ultrashield Av400 spectrometer, operating at 400.13 MHz, employing a 5 mm high-resolution broad-band ATMA gradients probe. Spectra were recorded using the zg30 pulse programme with P90 = 14.5 ␮s covering a sweep width 20.7 ppm (8278 Hz) with 64k time domain data points giving an acquisition time of 3.95 s, Fourier transformed using 128k data points and referenced to an internal TMS standard at 0.0 ppm. 2.7. Micellization and drug loading PEG-Hyd-PCL micelles, unloaded and loaded with ␣-tocopherol, were formed with the solvent-displacement method (Aliabadi et al., 2007). Briefly, 25 mg of the polymer were dissolved in 5 mL of either tetrahydrofuran, THF, or in an ␣-tocopherol solution in THF. The mass ratio of polymer/drug varied from 1:0.5 to 1:2. Then, using a peristaltic pump, the solution was transferred to 25 mL of deionised water with a flow rate of 0.5 mL/min. The suspension formed was gently stirred for 6 h in the dark, while the pH was monitored and constantly adjusted to 7.4. After complete removal of THF under vacuum, the volume was adjusted to 25 mL with distilled water and the suspension was transferred into a separation funnel; left standing for 1 h and the lower 20 mL were collected. The exact procedure was also followed for ␣-tocopherol solutions in THF without any polymer (blanks). 1 mL of each sample was freeze-dried and re-dissolved in methanol. This resulted to the disruption of the micelles and the release of the drug to the solution. The samples were then analyzed using HPLC.

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2.8. Dynamic light scattering (DLS) The particle size distribution of the micelles was determined by DLS using a Malvern Instruments Zetasizer 3000 HSa, operating with a HeNe laser (633 nm) at a scattering angle of 90◦ . The concentration of the micelles in the samples was 1 mg/mL and the temperature was 25 ◦ C. Each sample was measured in triplicate. The polydispersity PDI was calculated by using the DLS data and the formula PDI = (/)2 where  is the standard deviation and  the average diameter (nm) 2.9. Transmission electron microscopy (TEM) The morphology of the particles was examined using a JEOL JEM-2000EX transmission electron microscope, equipped with an Energy Dispersive Spectroscopy (EDS) Si(Li) detector. Samples (1 mg/mL) were deposited on a 400 mesh carbon film copper grid for 1 min, excess sample solution was removed, and left to dry in air before analysis. 2.10. High-performance liquid chromatography (HPLC) HPLC analysis was performed using an Agilent 1100 HPLC System and a Supelco Ascentis C18 column, 5 ␮, 150 mm × 4.6 mm. The mobile phase was methanol:water (97:3, v/v), and the analyte was detected at 292 nm. The flow rate of the mobile phase was 1.2 mL/min and the column temperature was 30 ◦ C. Quantification of ␣-tocopherol was based on a calibration curve created by diluting a stock solution of 10 mg/mL ␣-tocopherol in methanol to concentrations 10, 5, 2.5, 1, 0.5, 0.25 and 0.1 mg/mL. 2.11. pH sensitivity Micelles were formed as described in Section 2.7. After removal of the organic phase, a phosphate or citrate buffer solution was added to produce 1 mg/mL aqueous micelle dispersions of 10 mM buffer strength and varying pH. After further 24 h, particle size distribution and zeta potential were calculated by DLS, Section 2.8. The optical transmittance of varied pH micelle dispersions was also measured in 24 h, at 500 nm, using a Perkin Elmer UV-Vis spectrophotometer, model Lamda 35. After exposure to different pH for 48 h, 1 mL samples were freeze dried, dissolved in THF and analyzed with GPC, Section 2.4. The critical micelle concentration, CMC, of the PEG-HYD-PCL micelles after 24 h exposure at different pH was determined using pyrene as a fluorescence probe, according to the method described by Olea et al. (2011). Steady-state fluorescence spectra were recorded on a Spex FluoroMax-4 spectrofluorometer by exciting at 339 nm. The relative emission fluorescence intensity of 372 nm (I1) and 384 nm (I3) varied as a function of copolymer concentration. The CMC was estimated as the point where the value of I1/I3 decreases steeply due to the transfer of the pyrene from water to a less polar medium (micelle core).

Table 2 Number average molecular weight, Mn , polydispersity index, Mw /Mn , and hydrophilic fraction, f, as calculated by gel permeation chromatography. Polymer

Mn

Mw /Mn

f

PEG-CHO-5 PEG-Hyd-OH PEG-Hyd-PCL-A PEG-Hyd-PCL-B PEG-Hyd-PCL-C PEG-Hyd-PCL-D

6295 6273 8075 9110 10,310 11,853

1.14 1.14 1.17 1.21 1.24 1.30

– – 0.76 0.67 0.59 0.52

2-HEH. This purification is essential, since polymerization of ␧-CL can be also initiated from the hydroxyl-end of the small molecule. Two aldehyde-terminated PEGs were modified with 2-HEH: one with molecular weight, MW, of 2 kDa, PEG-CHO-2, and one with 5 kDa, PEG-CHO-5. The aim was to examine the effect of polymer MW to the micellar size and to the drug loading capacity. Ethanol is a good solvent for PEG at elevated temperatures but under the conditions used (−18 ◦ C) PEG-CHO-5 losses, in the purification step, were limited (˛ = 72%). However, for the lower-MW PEG, polymer losses were substantial. The yield of the reaction was only 11.25%, rendering PEG-CHO-2 unsuitable for further use in this work. For this reason polymerization of ␧-CL was carried out only for the higher-MW hydrazone derivative, PEG-Hyd-PCL-5. 3.2. Polymer characterization The MW of the polymers synthesized, as calculated by GPC, is shown in Table 2. All polymers synthesized had small polydispersity indexes, PDIs, and unimodal MW distributions. The latter is essentially confirming successful removal of 2-HEH from the hydrazone PEG derivative, since otherwise, concurrent ␧-CL polymerization from 2-HEH would have led to a product with a bimodal MW distribution. The hydrophilic fraction, f, is the mass fraction of the hydrophilic block to the total polymer mass and is what dictates the structure of the nanoparticles. For diblock amphiphilic copolymers, Discher and Eisenberg (2002) suggest that micelles are formed if f > 0.5; a condition that is satisfied by all four synthesized polymers. FT-IR spectra for PEG-CHO-5, PEG-Hyd-OH and PEG-Hyd-PCLC are presented in Fig. 3. All three materials show characteristic absorbances for PEG as the C O C etheric bond bending vibration at 1109 cm−1 and the absorbencies at 842 and 1333 cm−1 , attributed to PEG crystalline regions. For both the hydrazone derivative and the diblock copolymer there is a significant increase at the absorption at 1630 cm−1 which can attributed to stretching of

3. Results and discussion 3.1. Polymer synthesis Regarding the hydrazone formation reaction, 2-HEH, a highly viscous fluid, was found to be insoluble in the majority of popular, inert, organic solvents except dimethyl sulfoxide, DMSO, and ethanol. Ethanol was preferred over DMSO due to the lack of miscibility of the later with ethers, the most commonly used precipitation solvents for PEG. After isolation of PEG-Hyd-OH, the washing with ethanol step was introduced for the complete removal of

Fig. 3. FT-IR spectra for the aldehyde-terminated PEG (M = 5 kDa), PEG-CHO-5, the hydrazone derivative, PEG-Hyd-OH, and the diblock copolymer PEG-Hyd-PCL-C.

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Fig. 4. Polymer chemical structure and proton numbering.

Table 3 Proton assignments. Proton(s)

PEG-CHO PEG-Hyd-OH PEG-Hyd-PCL

Chemical shift, ␦ a

b

c

d

e

f

g

h

i

j

k

l

m

n

o

p

q

r

3.4 3.4 3.4

3.7 3.7 3.7

3.8 3.8 3.8

3.5 3.5 3.5

6.2 6.2 6.2

2.2 2.2 2.2

1.7 1.7 1.7

2.5 2.5 2.5

9.8 7.7 7.7

7.0 7.1

2.8 2.8

3.7 3.7

– 4.0

1.7

1.3

1.7

2.3

–

the C N bond of the hydrazone link. Finally, on the PEG-HYD-PCL-C spectra, new absorbances emerge of which the one at 1724 cm−1 is attributed to stretching of the esteric carbonyl, while the two at 2935 and 729 cm−1 are due to C H bond stretching in the PCL block. All absorbancies attributed to the PCL block increase in intensity from PEG-Hyd-PCL-A to PEG-Hyd-PCL-D, as the molecular weight of the hydrophobic block increases respectively (data not shown). Chemical structure for polymers and proton numbering for 1 H NMR characterization is shown in Fig. 4, while proton assignments are presented in Table 3. Of these, noteworthy is that the absorbance at 9.8 ␦ in the spectrum of PEG-CHO, which is due to aldehyde proton. This absorbance is absent in the spectrum of PEG-Hyd-OH, suggesting complete conversion to the hydrazone derivative. Moreover, both PEG-Hyd-OH and PEG-Hyd-PCL show new peaks at 7.0 and 7.5 ␦ that are attributed to the protons of the hydrazone bond. Finally, for all PEG-Hyd-PCL copolymers, the degree of polymerization, DP, of PCL was calculated using the equation: DPPCL = (A4.0 /2)/(A3.3 /3) = (A2.3 /2)/(A3.3 /3). Absorbances at 4.0 and 2.3 ␦ are due to protons in the PCL block, while the absorbance at 3.3 ␦ is due to the three protons in the methoxy terminal-group of PEG. This also allowed the calculation of the molecular weight for each polymer using NMR spectroscopy. The results are in relatively good agreement with those obtained using GPC and are presented in Table 4. 3.3. Micellization and drug loading The micelles were formed following the procedure described in Section 2.7. Micellization was also attempted by adding the aqueous to the organic phase (Chen et al., 2009) and by using the dialysis technique (Choi et al., 2006). However, both methods resulted in the formation of micelles with average sizes bigger than 500 nm, which is clearly not in the desirable nanoscale for

pharmaceutical applications. The increased size might be attributed to solvent being trapped in the micelle core leading to swelling of the polymer. Moreover, the use of phosphate buffer (pH 7.4, 10 mM) for the aqueous phase resulted to the formation of micelles with a “semi-open” structure (Fig. 5a). These micelles incorporated a substance that EDS analysis showed that contained significant amounts of phosphorous, indicating that phosphates may have been trapped within the micelle core. This is the reason why plain distilled water was used instead, while the pH was constantly adjusted to 7.4 in order to avoid hydrolysis of the acid-sensitive hydrazone bond. Micelles formed using distilled water, had a core–shell spherical structure (Fig. 5b). The particle size distributions of the PEG-Hyd-PCL micelles, unloaded and loaded with various ratios of ␣-tocopherol are shown in Fig. 6. The mean micelle diameters for increasing MW and increasing drug loading are shown in Fig. 7 and the polydispersity is shown in Table 5. The particle size distribution (PSD) and mean micelle sizes depended on the length of hydrophobic block, on the f ratio, and on the MW in agreement with the experimental studies of Riley and coworkers (Riley et al., 2001; Heald et al., 2002) who showed the dependence of the micelle size on the length of the constituent blocks, and also on the drug loading. The MW of the hydrophobic block in this work ranged approximately between 1800 and 5600 g/mol and increased from PEG-Hyd-PCL-A (A) to PEG-Hyd-PCL-D (D), while f decreased as shown in Table 4. Fig. 6 (No drug) shows the particle size distributions of the polymers for increasing length of hydrophobic block (or MW) from A to D, and decreasing f, respectively, in the absence of a-tocopherol. The PSD of polymer A with the smallest molecular weight was broader and shifted towards larger sizes. As the molecular weight increased from 8 K for polymer A to about 9.1 K for polymer B, the PSD shifted towards smaller sizes and became narrower. Micelles formed by B exhibited the smallest average particle size of 65 nm as shown in

Table 4 Degree of polymerization for PCL, polymer molecular weight as calculated by 1 H NMR and GPC, and CMC. Sample

DPPCL

Molecular weight 1

PEG-Hyd-PCL-A (A) PEG-Hyd-PCL-B (B) PEG-Hyd-PCL-C (C) PEG-Hyd-PCL-D (D)

14 25 42 65

H NMR

7828 9178 11407 13,687

CMC (␮g/mL) GPC

Variation (%)

8075 9110 10,310 11,853

-2.63 0.75 6.67 13.40

3.6 2 1.9 1.5

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Fig. 5. TEM images of PEG-Hyd-PCL-C micelles formed in (a) phosphate buffer and (b) distilled water.

Fig. 6. Particle size distribution for PEG-Hyd-PCL micelles with different ␣-tocopherol content.

Fig. 7. Further increase of the MW from B to C and D resulted in broader PSDs, which were shifted towards larger sizes. The results suggest that micelle size might not depend solely on the MW; other contributing factors may exist. It seems that f, apart from dictating the nano-structure, also had a significant effect on the PSD. The initial decrease of the particle size from A to B and its subsequent increase from B to C and D, with increasing length of the hydrophobic block, could be explained by the competing action of two co-existing phenomena. These two phenomena are the hydrophobic interactions that cause the micelles to form, and the length of the polymer blocks. Stronger hydrophobic interactions cause the polymer blocks to pack more tightly leading to smaller micelle sizes, whereas increasing the length of the hydrophobic block would lead to an increase in the MW and hence the diameter of the micelles (Hafezi and Sharif, 2012). Initially, the decrease

of the micelle size with increasing length of the hydrophobic block from A to B could be attributed to the stronger hydrophobic interactions due to the longer hydrophobic blocks that cause polymer B to pack more densely than polymer A, forming smaller micelles. In this case the effect of the hydrophobic interactions prevailed compared to the effect of the MW. When the length of the hydrophobic block and the MW increased further from B to C and D, the effect of the MW is likely to have prevailed leading to the formation of larger micelles. The dependence of the micelle size on the hydrophobic block length (NA ), could be expressed by the relationship Rhyd ˛NA 0.41 for polymers B, C and D, showing that the radius increased with increasing hydrophobic block length. This relationship differs than other power laws reported for star micelles (Letchford and Burt, 2007). Polymer A, with the smallest hydrophobic block did not follow the same trend as mentioned in the above paragraph,

Table 5 Mean diameters and polydispersity for all the polymers and drug loadings. Polymer

No drug

PEG-Hyd-PCL-A PEG-Hyd-PCL-B PEG-Hyd-PCL-C PEG-Hyd-PCL-D

122 65 88 95

Size (nm) ± ± ± ±

21.6 19.6 24.6 23.2

Polymer/drug 1:0.5 PDI

Size (nm)

0.032 0.09 0.062 0.144

124 126 136 142

± ± ± ±

3.5 4.3 4.5 4.8

Polymer/drug 1:1 PDI

Size (nm)

0.001 0.001 0.001 0.001

142 152 178 183

± ± ± ±

18.0 18.8 25.5 25.7

Polymer/drug 1:2 PDI

Size (nm)

0.016 0.015 0.020 0.020

166 178 185 185

± ± ± ±

24.7 30.2 28.4 28.2

PDI 0.022 0.029 0.023 0.023

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Fig. 7. Average particle size for PEG-Hyd-PCL micelles with different ␣-tocopherol content.

probably due to its increased solubility and hence to less dense packing of the core. However, when the micelles encapsulated ␣-tocopherol in their hydrophobic core, the average particle size increased monotonically, as the molecular weight increased. This phenomenon could be attributed to the fact that higher molecular weight equals to longer PCL blocks which could incorporate higher amounts of the drug leading to swelling of the micelles. This is supported from the fact that the HPLC analysis showed that ␣-tocopherol content increased from A to D micelles for all polymer/drug ratios (Table 6). When the hydrophobic drug is added at a small concentration, polymer/drug = 1:0.5 (Fig. 6), all particle size distributions become significantly narrower, the PDI decreases significantly, indicating that the hydrophobic drug may be enhancing the cohesion and the packing density of the hydrophobic part. Polymer A gives same size micelles with and without drug (Fig. 7), also indicating that the presence of the drug enhances the cohesion of the hydrophobic part, causing the polymer chains to pack more densely. Incorporation of ␣-tocopherol in liposomes has been found to decrease particle size (Bradford et al., 2003; Padamwar and Pokharkar, 2006). This has been attributed to the increased cohesion of the apolar, hydrophobic part of the membrane, which comes as a result of ␣-tocopherol inclusion. However, in accordance with ˜ et al. (2008), the present work, both Forrest et al. (2006) and Yánez who studied PEG-PCL micelles, reported increased particle sizes with increased ␣-tocopherol content. Interestingly, although the average particle size increased with the incorporation of ␣-tocopherol in the micelles, the width of the distribution decreased. This indicates that ␣-tocopherol might have been involved in the micellization process, most probably acting as a nucleation agent and giving a narrower size distribution. This effect was more profound for low ␣-tocopherol contents since

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swelling of the core seemed to mask it in higher drug contents. Furthermore, several researches have reported that core-entrapped drug, either chemically bound or physically entrapped, can lead to increased micelle stability. Kataoka et al. (2001) suggest that in these cases the drug acted as a filler molecule. PEG-Hyd-PCL micelles proved to be very efficient in encapsulating ␣-tocopherol as shown in Table 6. As explained earlier, encapsulation efficiency increased with increasing MW. For 1:0.5 polymer/drug ratio, almost all of the drug was encapsulated by the highest MW polymer, whereas, ever for the lowest MW polymer, efficiency remained above 85%. Polymers C and D efficiently encapsulated ␣-tocopherol at polymer/drug feed even as high as 1:1. The encapsulation efficiency was lower for the two other polymers, where almost 1/3 of the drug remained in the micellization medium. Further increase of drug loading in the feed resulted to the increase of the ␣-tocopherol content in the micelles, but the efficiency dropped for all polymers, except for D, which exhibited higher encapsulation efficiency in all three scenarios. Polymer D produced tocopherol-loaded micelles with the biggest average particle size. Nevertheless, the size remained below the 200 nm milestone.

3.4. pH-sensitivity and CMC The pH-sensitivity of the micelles originates from the pHsensitive hydrazone bond which is stable at neutral pH but dissociates at mildly acidic pH (Kalia and Raines, 2008). Transmittance measurements for polymer B for different pH values (Fig. 8(a)) showed that the transmittance decreased with decreasing pH as has been reported for pH sensitive polymers by other workers (Oh et al., 2009; Liu et al., 2012) indicating the dissociation of the micelles. Hydrolysis of the hydrazone bond led to the formation of PCL that is insoluble in the aqueous solution, causing the transmittance to decrease (Liu et al., 2012). The solution remained relatively transparent to light at pH 7.5 (77.5% transmittance), while the transmittance decreased significantly with decreasing pH, to as low as 22.5% at pH 5, suggesting that the hydrolysis rate increased with decreasing pH and that micelle dissociation occurred. After hydrolysis of the hydrozone bond, PEG polymer blocks dissolve in the solution, whereas the PCL blocks being insoluble tend to agglomerate forming larger particles. As a result the micelle size increased with increasing degree of hydrolysis or decreasing pH. As shown in Fig. 8(a), the micelle size changed slightly for pH values from 7.5 to about 6, but increased sharply for further pH decrease to 5.5 or lower. Fig. 8(b), shows that the size distribution of the micelles changed slightly from pH 7.5 to 6, whereas it changed significantly at pH 5 by shifting towards larger particle sizes that could be explained by the agglomeration of PCL blocks. This agglomeration at pH 5 eventually led to phase separation after 3 days. Phase separation was not observed at neutral

Fig. 8. (a) Transmittance (%) and micelle size for different pH (b) micelle size distributions for various pH values.

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Table 6 Encapsulation efficiency and drug content for different polymer/␣-tocopherol mass ratios. Sample

Polymer/drug 1:0.5

PEG-Hyd-PCL-A PEG-Hyd-PCL-B PEG-Hyd-PCL-C PEG-Hyd-PCL-D

Polymer/drug 1:1

Polymer/drug 1:2

Encaps. efficiency

Drug content (mg/mL)

Encaps. efficiency

Drug content (mg/mL)

Encaps. efficiency

Drug content (mg/mL)

0.86 0.91 0.96 0.99

0.43 0.45 0.48 0.5

0.62 0.70 0.83 0.93

0.62 0.70 0.83 0.93

0.37 0.43 0.65 0.80

0.74 0.86 1.30 1.60

Table 7 Transmittance (%) and zeta potential at different pH. %Transmittance

No tocopherol Polymer/tocopherol 1:1

Zeta (mV)

pH 7.5

pH 6

pH 5

pH 7.5

pH 6

pH 5

77.23 66.16

50.34 61.01

22.39 44.53

−11.5 −9.5

−9.8 −8.3

−5.5 −4.0

Fig. 9. Molecular weight distributions at different pH.

pH. A similar decreasing trend was found for the transmittance of solutions with tocopherol loaded micelles as shown in Table 7. The increase of particle size and reduced micelle stability with decreasing pH was also supported by molecular weight measurements carried out by GPC. Firstly, the molecular weight distributions (MWD) of the micelles measured at different pH values (Fig. 9) showed that the initial MWD of the polymer was mono-modal but it became bimodal with decreasing pH, showing that hydrolysis of the hydrazone bond and micelle dissociation occurred. The peak appearing at approximately 7 min, which is the first peak in the bimodal distribution at pH 6 and 5, corresponded to an average MW 9000, which was attributed to the

MW of the diblock copolymer, whereas the second peak appearing after 7.5 min, corresponded to an average MW 4700, which was attributed to the MW of either of the blocks. Also the ratio of the area or height of the second peak to the first increased as the pH decreased from pH 6 at pH 5 indicating that the hydrolysis rate is faster for more acidic pH. Zeta potential measurements (Table 7) were also in agreement with the previous findings, as they showed that the zeta potential decreased for increasing pH for both drug loaded and unloaded micelles. Decreasing zeta potential indicates reduced micelle stability at lower pH. The emission fluorescence spectra of pyrene in phosphate solutions (pH 7.4, 10 mM) of increasing polymer concentration (C = 0.1–500 ␮g/mL) for polymer B are shown in Fig. 10(a). The emission fluorescence spectra of pyrene in other buffered aqueous media of varied pH were also similar. Fig. 10(b) shows the plot of the ratio I372 /I384 , I1/I3, against concentration of block copolymer at pH 7.4 and 6. In low polymer concentration the ratio indicates that pyrene is exposed to a polar medium (water). The steep drop in the I1/I3 ratio when polymer concentration further increases, denotes the transfer of pyrene from water to a less polar medium (micelle core) and hence, the minimum polymer concentration to form micelles, i.e. critical micelle concentration, CMC. The CMC values at pH 7.5 and 6 for polymer B were found to correspond to the same polymer concentration, of approximately 2 ␮g/mL. The I1/I3 ratio for pH 5, did not show any trend, instead it showed very random and irregular behaviour, which could be attributed to the failure of micelles to form due to the hydrolysis of the hydrazone bond and the dissociation of the two blocks of the copolymer. The CMCs for all the polymers at pH 7.4 are shown in Table 4. The CMC is higher for polymer A due to the smaller hydrophobic block length which renders the polymer more soluble than the polymers with longer hydrophobic blocks. The CMC decreases with increasing length of the hydrophobic block as expected, due to the decreasing solubility. The increase of the hydrophobic block

Fig. 10. (a) Emission fluorescence spectra of pyrene in phosphate solutions (10 mM) at pH 7.4 for PEG-Hyd-PCL-B (b) I1/I3 versus concentration.

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