The influence of lipid composition and surface charge on biodistribution of intact liposomes releasing from hydrogel-embedded vesicles

International Journal of Pharmaceutics 459 (2014) 30–39

Contents lists available at ScienceDirect

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

The influence of lipid composition and surface charge on biodistribution of intact liposomes releasing from hydrogel-embedded vesicles A. Alinaghi a , M.R. Rouini a,∗ , F. Johari Daha b , H.R. Moghimi c a

Biopharmaceutics and Pharmacokinetic Division, Department of Pharmaceutics, Faculty of Pharmacy, Tehran University of Medical Sciences, Tehran, Iran Radioisotope Division, Nuclear Research Center, Atomic Energy Organization of Iran, Tehran, Iran c Department of Pharmaceutics, Faculty of Pharmacy, Shaheed Beheshti University of Medical Sciences, Tehran, Iran b

a r t i c l e

i n f o

Article history: Received 5 September 2013 Received in revised form 4 November 2013 Accepted 5 November 2013 Available online 13 November 2013 Keywords: Radiolabeled liposomes In situ forming liposomal hydrogel Liposome composition Liposome charge IP injection Biodistribution

a b s t r a c t Mixed drug delivery systems possess advantages over discrete systems, and can be used as a strategy to design more effective formulations. They are more valuable if the embedded particles perform well, rather than using drugs that have been affected by the surrounding vehicle. In order to address this concept, different liposomes have been incorporated into hydrogel to evaluate the potential effect on the controlled release of liposomes. Radiolabeled liposomes, with respect to different acyl chain lengths (DMPC, DPPC, or DSPC) and charges (neutral, negative [DSPG], or positive [DOTAP]) were integrated into chitosan-glycerophosphate. The results obtained from the biodistribution showed that the DSPC liposomes had the highest area under the curve (AUC) values, both in the blood (206.5%ID/g h−1 ) and peritoneum (622.3%ID/g h−1 ), when compared to the DPPC and DMPC formulations, whether in liposomal hydrogel or dispersion. Interesting results were observed in that the hydrogel could reverse the peritoneal retention of negatively charged liposomes, increasing to 8 times its AUC value, to attain the highest amount among all formulations. The interactions between the liposomes and chitosan-glycerophosphate, confirmed by the Fourier transform infrared (FTIR) spectra as shifted characteristic peaks, were observed in the combined systems. Overall, the hydrogel could control the release of intact liposomes, which could be manipulated by both the liposome type and interactions between the two vehicles. © 2013 Published by Elsevier B.V.

1. Introduction Combining biomaterials and technologies to design and develop advanced mixed drug delivery systems, and effective drug formulations, seems to be an attractive field for research. In this circumstance, numerous studies have been reported in the literature dealing with the integration of polymeric drug delivery systems and drug-loaded liposomes (Chung et al., 2006; Hara and Miyake, 2001; Ionov et al., 2011; Mulik et al., 2009; Stenekes et al., 2001). This combined system was developed to complement the advantages, while avoiding the disadvantages, of both the liposomal and polymeric systems. Moreover, the release of the encapsulated drug, and consequently, its pharmacokinetic parameters, could be modified through sustained drug action (Mulik et al., 2009).

∗ Corresponding author at: Biopharmaceutics and Pharmacokinetic Division, Department of Pharmaceutics, Faculty of Pharmacy, Tehran University of Medical Sciences, 14155-6451 Tehran, Iran. Tel.: +98 2166959056; fax: +98 2166959056. E-mail address: [email protected] (M.R. Rouini). 0378-5173/$ – see front matter © 2013 Published by Elsevier B.V. http://dx.doi.org/10.1016/j.ijpharm.2013.11.011

Liposomes represent one of the safest, most unique and potentially versatile vehicles used to date, with respect to a wide range hydrophilic or hydrophobic drug encapsulations, good biocompatibility, low toxicity, lack of immune system activation, and targeted delivery of drugs to the site of action (Chen et al., 2010; Immordino et al., 2006). Liposomal research has come a long way from its initial discovery, and now different liposomes are being engineered, in terms of size, lipid composition, or surface characteristics for specific applications (Banerjee, 2001). However, the application of conventional liposomes is mainly limited by their instability, short half-lives, and rapid clearance by the mononuclear phagocyte system (MPS). To circumvent these drawbacks, two polymeric approaches have been suggested so far: first, the surface modification of liposomes with hydrophilic polymers, such as polyethylene glycol (PEG), to produce MPS-evading vesicles; and second, the incorporation of liposomes within depot polymer-based systems (Mufamadi et al., 2011). The protection of liposomes from recognition by cells of the MPS is due to the steric hindrance effects of the polymers to opsonins (Boerman et al., 2000), while liposomal polymeric systems offer the possibility of a controlled release of intact liposomes from a

A. Alinaghi et al. / International Journal of Pharmaceutics 459 (2014) 30–39

reservoir in a sustained way (Alinaghi et al., 2013; Stenekes et al., 2001). The benefits of the latter composite system comprise an improvement in liposome stability, increased efficacy, the ability to control drug release for a longer period of time, reduced systemic toxicity, and the protection of sensitive drugs in a polymeric-based technology (Mufamadi et al., 2011). Moreover, local drug retention and sustained drug release for longer periods of time profit from creating an ideal reservoir for drugs or liposomes (Budai et al., 2007; Hurler et al., 2012; Mulik et al., 2009; Nie et al., 2011). In this regard, the in situ formation of hydrogels, because of their biodegradability, biocompatibility and nontoxicity, have been considered for delivering drugs directly to the site of action, through subcutaneous, intraperitoneal, or intratumoral injections. This highlights the valuable potential use of liposomal hydrogels as alternative strategies for topical, ocular, and chemotherapy, to improve both the efficacy and residence time of drugs in a targeted site. Ocular drug delivery benefits from liposomal hydrogels, since ocular therapy has long been a challenging task, mainly due the insufficient residence time of liquid formulations in the conjunctival sac (Sultana et al., 2006). The effects of different polymers for gel formation and lipid composition on the in vitro release of ciprofloxacin were studied by Budai et al. (2007). Polyvinyl alcohol and polymethacrylic acid (PMA) derivatives were applied in various concentrations in the presence of vesicles from lecithin and DPPC. Among the liposomal formulations, DPPC-PMA 0.1% presented the highest ciprofloxacin half-time, indicating that both the lipid composition with respect to the saturation of the phospholipid, and the type of polymer, were responsible for the differences found in the in vitro release of the antibacterial agent. Topical drug delivery is another area that the application of liposomal hydrogel as an alternative strategy has explored. As an example, mupirocin-in-liposomes-in-hydrogels were proposed as an advanced delivery system for improved burn therapy by Hurler et al. (2012). This formulation resulted in the prolonged release of liposomally associated mupirocin (both in vitro and ex vivo), remarkable bioadhesiveness, and antimicrobial potential effects. In addition, the drug release was affected by vesicle size, so that the hydrogel formulations containing smaller liposomes exhibited significantly lower release rates, when compared to the micron sized multilamellar vesicles. Obviously, liposomal hydrogels are promising carriers that may be loaded with a wide spectrum of chemotherapy drugs, to be used in parenteral formulations for treating local cancers. Nie et al. (2011) developed a thermosensitive Pluronic based hydrogel containing liposomes, for the controlled delivery of paclitaxel. Due to the increased viscosity of the liposomal gel, which has the effect of creating a drug reservoir, the longest drug-release period, compared with the liposome, gel, and the commercial formulation Taxol® , was observed. The majority of proposed mixed drug delivery systems have focused on formulations which aim to sustain the release of drugs over prolonged time. To the best of our knowledge, the current research could be differentiated from others in monitoring the escape of intact liposomes as a second vehicle from surrounded hydrogel in injection site and its fate in the body which could be affected by liposome formulation versatility. In a previous study, we focused on the release of intact radiolabeled liposomes from a chitosan ␤-glycerophosphate in situ forming hydrogel, and its tissue distribution after intraperitoneal injection in mice. The results showed that this system could prolong the release of liposomes in the peritoneum, and increase the durability of liposomes in the blood, compared to the liposomes or hydrogel (Alinaghi et al., 2013). It has been known for many years that the fate of liposomes in body and tissue distribution could be manipulated by liposome properties, including size, fatty acyl chain length of the phospholipids, head group compounds and charges, and the inclusion of

31

other lipids (like cholesterol) (Charrois and Allen, 2004). In the current study, the effects of liposome compositions and surface charges, as well as the possible effects of hydrogel on the peritoneal retention and tissue distribution of embedded liposomes, have been explored, following the intraperitoneal injection of radiolabeled formulations into mice. 2. Materials and methods 2.1. Materials 1,2-Distearoyl-sn-glycero-3-phosphocholine (DSPC), 1,2-dipalmitoyl-sn-glycero-3 phosphocholine (DPPC), 1,2-dimyristoyl(DMPC), and 1,2-dioleoylsn-glycero-3-phosphocholine sn-glycero-3-phosphoethanolamine (DOPE) were purchased from Lipoid GmbH (Germany), and 1,2-dioleoyl-3trimethylammonium-propane, chloride salt (DOTAP) and 1,2-distearoyl-sn-glycero-3-phosphoglycerol, sodium salt (DSPG) from Northern Lipids Inc. (Vancouver, Canada). Cholesterol (chol), glutathione (GSH), and ␤-glycerophosphate were provided by Sigma (Germany). ChitoClear® (chitosan) with medium viscosity (400 KD), and a 95% degree of deacetylation, was obtained from Primex (Iceland). Nuclepore® polycarbonate membranes, with pore sizes of 100 and 200 nm and a diameter of 25 mm, were purchased from Whatman Int. Ltd. (Singapore, Malaysia), and chloroform, methanol, ammonium thiocyanate, FeCl3 ·6H2 O, KCl, NaCl, Na2 HPO4 ·12H2 O, and KH2 PO4 from Merck (Darmstadt, Germany). Sephadex G-25 fine and Sephadex G-50 medium were provided by Fluka (Switzerland). HMPAO (hexamethylpropyleneamineoxime) kits containing 0.5 mg HMPAO and 5 ␮g SnCl2 ·2H2 O were prepared in-house. 99m Tc (as sodium pertechnetate; TcO4 − Na+ ), was eluted from a 99 Mo/99m Tc generator. 2.2. Preparation of liposomes Liposomes with different compositions (Table 1) (different fatty acyl chain lengths and charges) were prepared by using the thinfilm hydration method, which was described previously (Alinaghi et al., 2013). For three formulations testing fluidity, a 50:45:5 molar ratio of phospholipid (DSPC, DPPC, or DMPC):cholesterol:DOPE was chosen. In the other cases, charged lipids were added at a concentration of 15 mol%. The lipids were dissolved in a mixture of chloroform/methanol (2:1, v/v) and the solvents were removed using a rotary evaporator under reduced pressure, at a temperature of 10–15 ◦ C higher than the transition temperature of the main phospholipid. The thin layer of lipid was then hydrated in a solution of 100 mM glutathione in phosphate buffered saline (PBS) (pH = 7.4), so that the total lipid concentration was 40 mM. Multilamellar vesicles were extruded through 0.2 ␮m and 0.1 ␮m polycarbonate membrane filters at the same temperature. The removal of the unentrapped glutathione was performed by passage over a 1.1 cm × 23 cm Sephadex G-50 column. 2.3. Characterization of liposomes The phosphatidylcholine content of the liposomes was determined using the Stewart method (1980) and the Table 1 Types of liposomes and symbol used in the text. Symbol

Composition

Molar ratio

DSPC DPPC DMPC DSPG DOTAP

DSPC:Chol:DOPE DPPC:Chol:DOPE DMPC:Chol:DOPE DSPC:Chol:DSPG:DOPE DSPC:Chol:DSPG:DOPE

50:45:5 50:45:5 50:45:5 35:45:15:5 35:45:15:5

32

A. Alinaghi et al. / International Journal of Pharmaceutics 459 (2014) 30–39

phosphatidylcholine content showed a yield of more than 80% for all of the prepared formulations. The average particle size, size distribution and ␨-potential of the liposomes were measured at room temperature in PBS (pH 7.4) by using a Malvern Zetasizer Nano ZS (Malvern Instruments, London, England). 2.4. Liposome labeling procedure The liposomes were labeled using the method developed by Phillips et al. (1992) and optimized by Mirahmadi et al. (2008). A kit of the lipophilic chelator, HMPAO, was reconstituted with 1 ml of saline containing 10 mCi 99m Tc-pertechnetate. After 5 min, 99m TcHMPAO was added to a suspension of liposomes encapsulating the GSH, to a final concentration of 3 mCi/1 ␮mol of lipids. The mixture was incubated at room temperature for 30 min, followed by the removal of the free labels by passage over a 1.1 cm × 23 cm Sephadex G-25 column. Labeling efficiencies were evaluated by measuring the activity of the formulations before and after column purification, using a gamma well counter.

quadrant, using a 21 G × 1½ in. (0.4 mm × 12 mm) syringe. The animals (three mice for each time point) were euthanized with carbon dioxide at 0.083, 0.25, 0.5, 1, 2, 4, 6, 12 and 24 h after injection. The mice were then weighed, and their blood collected by cardiac puncture. The ventral skin was carefully and completely removed, taking care not to damage the parietal peritoneum. Subsequently, normal saline (3 ml) was flushed into the peritoneal cavity and re-collected (Mirahmadi et al., 2010). Finally, the mice were dissected, and the liver, spleen, stomach, intestines, kidneys, heart, lungs and carcass were washed with normal saline, and blotted dry using Whatman filter paper. The radioactivity of each organ was measured using a gamma well counter after weighing. The amount of radioactivity of the organs of the mice was calculated using the radioactivity of the original formulations, at the same times, as references. The results were expressed as a percentage of the injected dose per gram of tissue (%ID/g), and %ID for the fluid collected from the peritoneal cavity. All mean values are given ±standard deviation (SD). The statistical comparisons were made by one-way ANOVA using SPSS 11.5 for Windows. The probability significance threshold was p < 0.05.

2.5. In vitro stability of radiolabeled liposomes 3. Results In our previous study, there was good stability with the DSPC liposomes when 0.3 mci/␮mol of radioactive lipid was added to the suspension (Alinaghi et al., 2013). Therefore, in this current research, an in vitro stability experiment was performed for the other vesicles, including DPPC, DMPC, DSPG and DOTAP. Radiolabeled liposomes (1 ml) were incubated with 3 ml of human plasma in a 37 ◦ C water bath, and 100 ␮l samples were withdrawn at 0.5, 1, 2, 4 and 24 h after incubation. The samples were then purified by passage through a 1.5 cm × 7 cm Sephadex G-25 column, and the results were compared to those before purification.

3.1. Characterization of liposomes The results from the particle size analysis and zeta potential measurements of the liposomes are shown in Fig. 1. For all liposomes, the size was uniformly distributed so that the polydispersity index was less than 0.2. The neutral liposomes (DSPC, DPPC, and DMPC) showed negative surface charges via anion adsorption (Li and Tian, 2006), and the inclusion of 15% charged phospholipids in the liposome changed the liposome charges to −22 and 15 mV for the DSPG and DOTAP, respectively.

2.6. Preparation of the chitosan-glycerophosphate solution 3.2. In vitro stability studies

In order to determine the interaction between the charged liposomes and chitosan hydrogel, an FTIR analysis of the formulations was performed. All formulations were lyophilized for 48 h using Christ Alpha 2–4 LD plus (Germany) freeze drier. The FTIR spectra were recorded with KBr pellets on a Magna-IR spectrometer 550 (Nicolet, USA). Prior analysis, transparent sample/KBr discs (ratio 1:10) were prepared by uniaxial pressing. The transmission spectra were recorded in the spectral range of 4000–400 cm−1 , 4 cm−1 resolution, at ambient temperature. 2.8. Animal experiments Animal experiments were carried out according to the Tehran University animal Care and Use guidelines. For this study, female NMRI mice, weighing between 20 and 35 g, were used, with access to water and mouse chow during the experiments. On the day of experiment, 1 ml of label free solution, label free hydrogel, radiolabeled liposomal dispersion, or hydrogel containing radiolabeled liposomes was injected intraperitoneally (IP) into the left lower

125

0.19

20 15

0.08

120

10 5

0.13

115

0 0.11

-5

0.17

110

-10 -15

105

ζ-potential (mV)

2.7. FTIR analysis

The results obtained from the in vitro stability of the liposomes after incubation in human plasma are shown in Table 2. The labeling efficiency in all of the experiments was acceptable at time zero, after that, releasing 99m Tc from the vesicles occurred in different amounts. Changes in the bilayer fluidity by reducing the acyl chain length did not considerably affect the stability of the liposomes. Less than 3% of the 99m Tc was released from the DSPC, DPPC, and DMPC at 30 min, and after that, the decline in the remaining radioactivity was only about 10% for each of them. Similarly, the positively charged liposomes showed desirable efficiency during the first hours, and the release of 99m Tc was 15% at 24 h post incubation. However, for the DSPG liposomes, the 99m Tc released very

Particle size (nm)

The hydrogel was prepared as outlined previously (Alinaghi et al., 2013). The chitosan was dissolved in a 0.1 M hydrochloric acid solution under stirring in an ice bath to prepare a 1.8% (w/v) solution, after which, ice-cold ␤-glycerophosphate aqueous solution (10%, w/v) was added dropwise to the iced chitosan-solution, and stirred to form a clear solution (sol phase). The pH value of the prepared sol phase was 7.3. Radiolabeled liposomes (3 ml) were added to the chitosan-glycerophosphate sol (4 g), and stirred gently for 10 min for uniform distribution of the liposomes in the solution.

-20 100

-25 DSPC

DPPC

DMPC

DSPG

DOTAP

Fig. 1. Average particle sizes (bars, left axis) with inserted polydispersity index (PdI) in figure, and zeta potential (dots, right axis) of different liposomes.

A. Alinaghi et al. / International Journal of Pharmaceutics 459 (2014) 30–39

33

Table 2 % Remained lable in different liposomes after incubation in human plasma at 37 ◦ C (mean ± SD, n = 3). Incubation time (h)

% Remained radioactivity DSPC

0 0.5 1 2 4 24

94.08 91.54 88.31 79.48 78.25 80.74

DPPC ± ± ± ± ± ±

1.79 2.29 2.43 2.93 2.44 2.61

84.34 81.41 73.71 82.57 77.68 73.95

DMPC ± ± ± ± ± ±

5.71 3.25 2.49 2.18 2.71 1.43

rapidly from the particles in such a way that 62% of the radiolabel remained at 30 min after incubation in the liposomes, and no further decrease was seen in the later times. 3.3. FTIR spectra In order to establish the interactions between the liposomes and chitosan hydrogel, FTIR spectra were collected. Freeze-dried liposomes (DSPC, DSPG, and DOTAP), the relevant liposomal hydrogel, and the chitosan-glycerophosphate hydrogel were analyzed for characteristic peaks (Fig. 2, first line in each spectrum). For all liposomal dispersions, the peak at around 1740 cm−1 represents the C O stretch of the ester bond, where the fatty acid chain meets the head group (Jiang et al., 2005). The CH2 symmetric and antisymmetric stretches (2800–3000 cm−1 ) and the absorption band (particularly at 3420, 3400, 3478 cm−1 for DSPC, DSPG, and DOTAP) appeared because of OH and NH stretching. For the chitosan-glycerophosphate hydrogel (Fig. 2, second line in each spectrum) several characteristic absorption bands were identified, such as the broad band in the region of 3000–3600 cm−1 resulting from the overlap of the OH and NH stretching; a smaller band at 2933 due to CH stretching; C O stretching at 1670 cm−1 ; NH in the primary amine and secondary amide groups at 1465 cm−1 ; C O C stretching at 1065 cm−1 ; and several peaks in the 600–500 region (bending motion of the O C N groups). The band at 528 arose from the P O asymmetrical bending of the PO4 −3 molecules (Douglas et al., 2013). The results obtained from each blend were interesting (Fig. 2, third line in each spectrum), since the hydrogel and liposome interactions were confirmed in the FTIR spectra recorded, with a shift of C O from 1740 cm−1 in the liposomes to 1661 cm−1 for the DSPC, 1721 cm−1 for the DSPG, and 1657 cm−1 for the DOTAP. Another interaction regarding the chitosan hydrogel spectrum was observed with a shift from 3364 cm−1 (regions of OH and NH in the hydrogel) to 3359, 3241, and 3391 cm−1 in the liposomal hydrogel, with the addition of DSPC, DSPG, and DOTAP.

88.47 90.59 87.23 85.75 75.32 78.85

DSPG ± ± ± ± ± ±

2.63 2.43 2.20 3.13 8.96 7.41

91.43 62.91 65.44 60.37 56.25 64.39

DOTAP ± ± ± ± ± ±

3.29 2.87 2.35 3.02 4.34 2.66

90.31 91.76 88.54 79.18 78.25 76.19

± ± ± ± ± ±

1.79 2.29 2.43 2.93 2.44 1.98

Peritoneal retention: Fig. 3 illustrates the percentages of the injected dose (%ID) at different times for the peritoneal lavage fluid. The AUC values for the radiolabeled liposomes and liposomal hydrogels accompanying their control groups, as well as the enhancement ratio, are represented in Tables 3 and 4. The enhancement ratio which signified the hydrogel effect on the peritoneal retention and tissue distribution of the liposomes was calculated from the division of the AUC values of the liposomal hydrogel, to the AUC values of the liposomal dispersion. As shown in Fig. 3, the non-particulate 99m Tc-HMPAO serving as the control group cleared

3.4. In vivo studies In the following two separate sections, the results of the intraperitoneal injection of the radiolabeled liposomes and liposomal hydrogel are presented. The effects of the different phospholipid compositions of the vesicles, followed by the effects of the liposome charge (considering the peritoneal level and biodistribution of the liposomes) are described. All statistical comparisons refer to the area under the curve (AUC). 3.4.1. Effects of lipid composition The results obtained after the intraperitoneal injection of the 99m Tc-HMPAO solution and hydrogel (control groups), and both liposomal formulations with different compositions (DSPC, DPPC, or DMPC:Chol:DOPE; Table 1) in the peritoneal cavity and other tissues, are depicted separately as follows:

Fig. 2. FTIR spectra for (A) DSPC, (B) DSPG, (C) DOTAP formulations. In each group first lines represent liposome dispersion, second lines chitosan-glycerophosphate, and the third black lines represent liposomal hydrogel.

34

A. Alinaghi et al. / International Journal of Pharmaceutics 459 (2014) 30–39

Table 3 Area under the %ID/g (or %ID) vs. time curves in various tissues for liposomal formulations with different compositions and control groups (AUC (SD)). Organ

Blood Peritoneal Wash Liver Spleen Intestines Stomach Kidneys Heart Lungs a b c d e

Liposome

Liposomal hydrogel

DSPC

DPPC

DMPC

99m

53.03a,e (11.75) 183.35a,e (35.82) 485.80a (67.57) 272.03a,c (34.14)d 44.41a,d (5.45) 29.62 (4.99) 52.85 (8.66) 4.97 (1.51) 3.61 (1.20)

56.06a,e (14.66) 174.5a (27.74) 489.99a (57.99) 189.51a,d (30.59) 11.37a,d (2.47)e 30.11 5.34 15.87 (2.23) 3.22 (0.70) 4.56 (0.97)

41.87a,e (7.28) 151.84a,e (17.94) 397.05a (27.70) 208.64a,c (12.20) 23.05a,b (2.09) 42.60 3.47 18.46 (3.15) 4.17 (1.13) 3.72 (0.85)

4.75 (1.39) 20.74 (3.32) 154.62 (19.76) 21.73 (4.26) 116.59 (3.65) 45.18 (3.74) 42.23 (4.29) 3.75 (0.61) 3.16 (0.75)

Tc-HMPAO

DSPC

DPPC

DMPC

99m

206.47a (29.75) 622.29a,c (92.28)d,e 544.44a,c (37.04) 201.28a (27.34) 51.07a (6.28) 64.22 (4.8) 64.66 (7.77) 4.73 (1.04) 3.29 (0.67)

189.25a (22.28) 307.36a,d (56.75)e 411.44a,b (57.73) 158.01a (30.35) 54.91a,e (2.38) 60.61 4.30 57.36 (3.24) 4.63 (0.86) 5.42 (1.02)

140.17a (28.06) 308.93a,c (31.86)e 312.48a,b (26.10)c 169.28a (29.62) 42.65a (4.17) 47.99 3.11 70.13 (2.91) 3.88 (0.49) 4.65 (0.63)

4.66 (0.92) 28.42 (5.70) 157.66 (12.99) 18.16 (5.89) 106.4 (6.61) 45.31 (6.93) 58.03 (3.99) 3.96 (0.40) 3.52 (0.95)

Tc-HMPAO

Formulations compared to relevant control. DPPC compared to DMPC. DSPC compared to DMPC. DSPC compared to DPPC. Each liposomes compared to respective liposomal hydrogel.

rapidly due to its low molecular weight and rapid absorption via the whole surface of the peritoneum (Flessner et al., 1984). Whereas all of the liposomal formulations, either dispersion or hydrogel forms, remained for a longer duration in significantly higher levels in the peritoneum (p < 0.05). The parallel peritoneal profiles, and the relatively similar AUC values of the liposomes with different compositions (Fig. 3a and Table 3), could suggest that the type of phospholipid has no effect on the retention of the liposomes in the peritoneum. However, the incorporation of the liposomes in the hydrogel, regardless of lipid type, increased the peritoneal retention when compared to the liposomal dispersion. For example, the AUC for the DSPC liposomal hydrogel was 3.4 times greater than the AUC for the DSPC liposomes (Table 4). Considering Fig. 3b, all formulations showed a similar radioactivity level in the first hours, with the differences appearing at alternate times, so that at 4 h post injection of the DSPC liposomal hydrogel, the radioactivity counts of the peritoneal lavage were nearly 2 times higher than those of the DPPC or DMPC activity (Fig. 3b). Profile in blood: As shown in Fig. 4, the incorporation of the liposomes in the hydrogel caused increasingly detectable radioactivity in the blood, with respect to both the level and duration, compared to the liposome dispersion and free lable (p < 0.05). Considering Fig. 4a, the percentage of the injected dose per gram of blood dropped to about 1% at 12 h for all liposomes, while the liposomal hydrogel (Fig. 4b) presented at 9.9, 9.1 and 6.5%ID/g of the DSPC, DPPC and DMPC liposomes in the blood at the same times, respectively. In our previous study, the effect of the gelation process on

introducing liposomes to the blood was described completely for the DSPC liposomes (Alinaghi et al., 2013). In the current experiments, similar profiles, including the ascending and descending trends, possessed two maximum points for the different phospholipids (Fig. 4b). These distinctive profiles verified that during the first hours, the liposomes left the dispersal media easily, due to the sol state, while the releasing liposomes was restricted after the gel formation. The liposomes could easily escape the cavity, and at 1 h the first maximum activity was observed. After that, the activity level in the blood decreased to nearly 5%ID/g at 2 h because of the gel formation. From 2 to 12 h, the radioactivity increased slightly, then decreased until 24 h post injection. However, the AUC of the DMPC dispersion and hydrogel formulations in the blood was lower than that of the DPPC and DSPC (Table 3), possibly due to the low transition temperature of the phospholipid (23 ◦ C), and more rapid uptake by the RES (Drummond et al., 1999). Distribution in other organs: The AUC values obtained from the %ID/g of various organs vs. time for the different formulations are listed in Table 3. All liposomes, in both the dispersion and hydrogel injected forms, concentrated significantly in the liver and spleen when compared to the label free (Table 3). Since 99m Tc-HMPAO is principally eliminated by the hepatobiliary pathway (Neirinckx et al., 1987), the liver showed a considerable AUC. In addition, the results showed that the 99m Tc-HMPAO, which is excreted in the GI tract (Neirinckx et al., 1987), had the highest level in both the intestines and stomach in the control groups. With respect to the enhancement ratios (Table 4), it seems that the increasing retention

Table 4 Enhancement ratio obtained from division of AUC values of liposomal hydrogel to AUC values of liposomal dispersion. Organ

DSPC

DPPC

DMPC

DSPG

DOTAP

99m

Blood Peritoneal Liver Spleen Intestines Stomach Kidneys Heart Lungs

3.89 3.39 1.12 0.74 1.15 2.17 1.22 0.95 0.91

3.38 1.76 0.84 0.83 4.83 2.01 3.61 1.44 1.19

3.35 2.03 0.79 0.81 1.85 1.13 3.80 0.93 1.25

3.06 8.18 0.63 0.94 8.08 1.93 3.44 0.65 0.86

6.24 1.02 0.90 1.70 1.75 0.67 3.63 1.12 0.76

0.98 1.37 1.02 0.86 0.91 1.00 1.37 1.06 1.11

a

Hydrogel divided to solution form.

Tc-HMPAOa

A. Alinaghi et al. / International Journal of Pharmaceutics 459 (2014) 30–39

a)

a) 100 DSPC Liposome

30

DSPC Liposome DPPC Liposome

25

DPPC Liposome

80

DMPC Liposome

DMPC Liposome 60

20

Tc-HMPAO Solution

Tc-HMPAO Solution

% ID/g

% ID

35

40

15 10

20

5 0 0

2

4

6

8

10

12

14

16

18

20

22

0

Time (h)

b)

0

24

b)

100

2

4

6

8

10 12 14 Time (h)

DMPC Liposomal hydrogel

22

24

DPPC Liposomal hydrogel DMPC Liposomal hydrogel

20

Tc-HMPAO Hydrogel

% ID/g

% ID

20

DSPC Liposomal hydrogel 25

DPPC Liposomal hydrogel

60

18

30

DSPC Liposomal hydrogel 80

16

40

Tc-HMPAO Hydrogel

15 10

20

5

0 0

2

4

6

8

10

12

14

16

18

20

22

24

Time (h) Fig. 3. Detectable radioactivity (in terms of %ID ± SD) vs. time in the peritoneal cavity with different compositions and free lable: (a) liposomes and (b) liposomal hydrogel.

time of the liposomes in the peritoneum resulted in an augmentation of the radioactive agent in the GI tract, while, except for the DSPC, there was no further surging effect on the liver and spleen uptake. In other words, among all formulations, the DSPC liposomes showed the highest AUC values in the liver and spleen, whether in the dispersion or hydrogel forms. One possibility for this is that during the residency of the liposomal hydrogel in the abdominal cavity, 99m Tc-HMPAO leakage from the aqueous phase of the DPPC and DMPC occurred, due to the lower phase transition temperature, although the in vitro test demonstrated that they were stable in human plasma at 37 ◦ C (Table 2). The AUC values for the kidneys, lungs and heart were calculated, and are listed in Table 3. There were no considerable levels of the liposomes or 99m Tc-HMPAO in the lungs or heart. 3.4.2. Effects of liposome charge The peritoneal retention and tissue distribution of the neutral (DSPC), negative (DSPG) and positive (DOTAP) vesicles (Table 5) were compared. Peritoneal retention: The percentage of the injected dose obtained from the peritoneal lavage fluid for the differently charged liposomes and label free are illustrated in Fig. 5, and the AUC values and enhancement ratios are shown in Tables 4 and 5. There were higher amounts of all vesicular formulations in the peritoneum, compared to the control groups (p < 0.05), except for the DSPG liposomal dispersion (p < 0.301). With regard to Fig. 5a, the incorporation of positively (DOTAP) and negatively (DSPG) charged lipids into the conventional vesicles had the opposite results when

0 0

2

4

6

8

10

12 14 16 18 20 22 24

Time (h) Fig. 4. Detectable radioactivity per gram (in terms of %ID/g ± SD) vs. time in blood for different liposome compositions and free lable: (a) liposomes and (b) liposomal hydrogel.

compared to the natural liposomes. The DOTAP increased the peritoneal retention of the liposome dispersion, while for the negatively charged liposomes, the peritoneal profile was completely different, and the incorporation of DSPG into these vesicles decreased the peritoneal retention. Interestingly, the injection of the liposomes in the hydrogel forms inverted this, so that the negative liposomes were retained in the abdominal cavity longer than the neutrally and positively charged ones. In other words, while the AUC values for the DOTAP liposomal hydrogel remained stable, compared to the liposomal dispersion, the enhancement ratios were 8.2 and 3.4 for the DSPG and DSPC, respectively. Profile in blood: As shown in Fig. 6, the absorption phase of the 99m Tc-HMPAO into the blood circulation could not be detected. Because of the low molecular weight and lipophilicity of 99m TcHMPAO, a rapid distribution into other organs would result in low blood activity. Almost all vesicles reached their maximum blood concentration 0.5 h after the IP injection of the liposomal dispersion, and 1 h post injection of the liposomal hydrogel. However, as seen in Fig. 6a, the DOTAP liposome dispersion had a very low level in the blood. In contrast, the DSPG showed the highest level; this behavior could be described by the peritoneal profile where the DSPG left the cavity more rapidly (Fig. 5a). The incorporation of the liposomes into the hydrogel caused several changes in the parameters (Fig. 6b). First, the maximum level increased to 1 h due to the gelation time in the mouse peritoneum, and at 2 h, all formulations experienced a trough. After that, the DSPC had a slight increase in the %ID/g, while both the DSPG and DOTAP remained almost stable

36

A. Alinaghi et al. / International Journal of Pharmaceutics 459 (2014) 30–39

Table 5 Area under the %ID/g (or %ID) vs. time curves in various tissues for differently charged liposomes and control groups (AUC (SD)). Organ

Liposome

Blood Peritoneal Wash Liver Spleen Intestines Stomach Kidneys Heart Lungs a b c d e

Liposomal hydrogel

DSPC

DSPG

DOTAP

99m

53.03a,d (11.75)e 183.35a,c (35.82)d,e 485.80a,d (67.57) 272.03a,c (34.14) 44.41a,c (5.45)d 29.62e (4.99) 52.85c,d (8.66) 4.97 (1.51) 3.61 (1.20)

39.48a,b (9.19)e 80.89a,b (28.76)c,e 435.88a,b (32.86)e 145.63a,c (20.76) 10.69a,b (1.52)c,e 23.30a,e (5.26) 8.32a,c (6.04) 5.36 (1.37) 5.08 (1.62)

15.68b,d (2.90)e 292.97a,b (40.06)d 244.15d (27.29) 191.43a (21.68) 21.84a,b (3.82)d 83.46a,e (5.37) 9.83a,d (5.38) 4.13 (1.19) 20.44 (3.02)

4.75 (1.39) 20.74 (3.32) 154.62 (19.76) 21.73 (4.26) 116.59 (3.65) 45.18 (3.74) 42.23 (4.29) 3.75 (0.61) 3.16 (0.75)

Tc-HMPAO

DSPC

DSPG

DOTAP

99m

206.47a,c (29.75)d,e 622.29a,d (92.28) 544.44a,c (37.04)d 201.28a (27.34) 51.07a (6.28) 64.22a,e (4.8) 64.66d (7.77) 4.73 (1.04) 3.29 (0.67)

120.65a,c (19.22) 662.01a,b (62.94) 273.76a,c (18.14) 137.86a,b (13.67) 86.38e (11.04) 44.94e (5.07) 28.64a,c (2.3) 3.51 (0.93) 4.38 (1.12)

97.79a,d (16.98) 300.38a,b (45.73)d 271.75d (34.52) 325.86a,b (40.83) 38.33a (12.53) 56.30e (7.08) 35.73d (2.14) 4.64 (0.97) 15.47 (2.81)

4.66 (0.92) 28.42 (5.70) 157.66 (12.99) 18.16 (5.89) 106.4 (6.61) 45.31 (6.93) 58.03 (3.99) 3.96 (0.40) 3.52 (0.95)

Tc-HMPAO

Formulations compared to relevant control. DSPG compared to DOTAP. DSPC compared to DSPG. DSPC compared to DOTAP. Each liposomes compared to relevant lipogel.

a)

a)

30

100

DSPC Liposome

DSPC Liposome DSPG Liposome

80

20

Tc-HMPAO Solution

60

10

20

5

0

0 0

2

4

6

8

10

12

14

16

18

20

22

24

Time (h)

0

2

4

6

8

b)

10 12 14 16 18 20 22 24 Time (h)

30

100

DSPC liposomal hydrogel

DSPC liposomal hydrogel

25

DSPG Liposomal hydrogel

80

20 % ID/g

Tc-HMPAO hydrogel

60 40

DSPG Liposomal hydrogel DOTAP Liposmal hydrogel

DOTAP Liposomal hydrogel % ID

Tc-HMPAO Solution

15

40

b)

DSPG Liposome DOTAP Liposome

DOTAP Liposome

% ID/g

% ID

25

Tc-HMPAO hydrogel

15 10

20

5

0 0

2

4

6

8

10

12 14 Time (h)

16

18

20

22

24

Fig. 5. Detectable radioactivity (in terms of %ID ± SD) vs. time in the peritoneal cavity for differently charged liposomes and free lable: (a) liposomes and (b) liposomal hydrogel.

0 0

2

4

6

8

10 12 14 Time (h)

16

18

20

22

24

Fig. 6. Detectable radioactivity per gram (in terms of %ID/g ± SD) vs. time in blood for differently charged liposomes and free lable: (a) liposomes and (b) liposomal hydrogel.

A. Alinaghi et al. / International Journal of Pharmaceutics 459 (2014) 30–39

from 2 to 12 h post injection. Second, the enhancement ratios were 3.9, 3.1 and 6.2 for the DSPC, DSPG and DOTAP, respectively, which means that the hydrogel had the greatest effect on the DOTAP formulation. Third, while all liposome dispersions were not detectable at 12 h, using the hydrogel form of the liposomes increased the presence of vesicles in the blood for 24 h. Distribution in other organs: The area under the %ID/g vs. time curves for the free lable and particulate systems in different organs are depicted in Table 5. The AUC values of the RES (liver and spleen), whether in the control groups or liposomes were considerable. The liver uptake of the different liposomes was a little complicated. For example, although the DOTAP liposomes could increase the liver uptake when compared to the control group, this increment was not significant (p < 0.06 and p < 0.27 for the dispersion and hydrogel, respectively). A similar result was observed for the DSPG liver uptake, as injected in the hydrogel form into the mice (p < 0.06). In addition, the accumulation of liposomes in the spleen increased significantly, ranging from 7.6 for the DSPG liposomal hydrogel to 17.9 for the DOTAP, when compared to the control groups. For other organs, like the intestines, stomach and kidneys, it seemed that the increased AUC values after using the liposomal hydrogel may be due to the increasing peritoneal retention of the formulations. Moreover, the DOTAP liposomes tended to accumulate in the lungs, and the AUC increased to 6.5 and 4.4 times for the liposomal dispersion and hydrogel, compared to the control groups, respectively. 4. Discussion Introducing a combined system of embedded particles, which allows intact liposomes to release from a hydrogel vehicle, was the first aim of our previous study (Alinaghi et al., 2013). This system is a promising strategy to control drug release in two steps by providing a vehicle to retain the particles, and release them for a longer period of time, followed by the drug release. In addition, the advantages of local drug delivery can be achieved through the injection of an in situ gel form into the target site. The present study is an effort to further investigate the effects of hydrogel on the release and disposition of liposomes, with different phospholipid compositions and charges followed in the dispersion and hydrogel forms, after being injected IP into mice. The peritoneal levels of all of the vesicular formulations were much higher than the free label formulations, which was consistent with previous reports (Chen et al., 2007; Dadashzadeh et al., 2010; Sadzuka et al., 2000). The relatively large well blood perfused surface area of the peritoneum provides an excellent site for drug exchange with the plasma (Flessner et al., 1984). However, two possibly different pathways of peritoneal absorption create the differences between small and particulate material levels in the abdominal cavity. Drugs less than 1000 Da are absorbed so rapidly from the whole surface of the peritoneum that most drain into the portal cavity (Howell, 2008). In contrast, particles larger than 20 nm, or compounds greater than 20 kDa, are absorbed mainly via the lymphatic ducts, which are located on the mouse subdiaphragm surface, with a mean opening diameter of 3.6 ± 2.0 ␮m (Tsai et al., 2007). In spite of the different transition temperatures of the DSPC (53 ◦ C), DPPC (42 ◦ C with a pretransition at 37 ◦ C), and DMPC (23 ◦ C) included in each liposome, a similar profile of the percentage of the injected dose vs. time in the peritoneal cavity for the three types of phospholipids was obtained. One possibility for this was the short residence time of the formulations in the site of injection, so there was no opportunity to measure the effects of the liposome compositions. This was confirmed when the liposomal hydrogels were tested. Applying hydrogel as a second vehicle, with higher viscosity compared to the liposomal dispersion, restricted the motility of

37

the vesicles in the first hours, and via gel formation in the later times. In addition, the release of the liposomes was slower due to the interaction between the chitosan and liposomes (Filipovicgrcic et al., 2001), and the formation of the hydrogen bonding between the polysaccharides and phospholipid head groups (Guo et al., 2003). Therefore, the retainment duration of the liposomes in the peritoneum increased, and differences between the formulations appeared. In these points, the liposomes containing the DSPC had the highest levels, which were reflected in the respective AUC values. These results are in accordance with a study reported by Dadashzadeh et al. (2010), where there was no significant difference between the different liposome compositions with the 100 nm particle size, after IP injection, while increasing the particle size to 1000 nm prolonged the retention time via the sedimentation of the vesicles. This caused a slight surge in the DSPC peritoneal AUC when compared to the other phospholipids. The liposomal hydrogel provided sustained levels, not only locally but also systemically, for a relatively long time (Fig. 4b). There was no significant difference between the AUC blood values of the liposomes with different compositions, whether in the liposomes or liposomal hydrogel. Conversely, the slow release of the vesicles from the chitosan-glycerophosphate hydrogel, followed by the liposomes entering the blood, resulted in a longer duration of the liposomes in circulation. It should be mentioned that a small increase in the AUC values of the DSPC liposomes could be related to a higher transition temperature of the phospholipids, creating a more rigid membrane bilayer. This type of liposome resistant penetration by the serum opsonins, and macrophage uptake, was diminished (Drummond et al., 1999). On the other hand, this rigidity may also decrease the 99m Tc-HMPAO leakage rate (Dadashzadeh et al., 2010). However, in vitro studies have shown that both the DPPC and DMPC vesicles are stable. Increasing peritoneal retention has a beneficial impact on the treatment of cancers originating in the peritoneal cavity, such as gastrointestinal, ovarian, and colorectal (Flessner, 2007). In this respect, Zahedi et al. (2009) presented a chitosan-phospholipid blend to localize the delivery of docetaxel following IP administration. Their product was found to result in a sustained plasma concentration and constant levels of the drug in the peritoneal cavities of healthy mice over 2 weeks. In these experiments, the tissue distribution of the liposomes was not affected by phospholipids with different acyl chain lengths. Pharmacokinetic comparisons were previously made across a series of different liposomes, including DMPC, DSPC, DOPC, POPC, and SMPC, in combination with cholesterol (2:1 molar ratio), by Charrois and Allen (2004). They showed that the accumulation of liposomal lipids into given tissues was not dependent on the formulation, whereas phospholipids with different acyl chain lengths or degrees of saturation could alter the liposome fluidity and drug release, so that the accumulation of the drug in these tissues could be affected. The surface potential is an important factor that has been found to influence the distribution of liposomes. In charged liposomal dispersion, positively charged liposomes showed the highest peritoneal levels, due to the electrostatic interactions between the positively charged liposomes and the negative surface of the peritoneal mesothelium, as well as the low tendency of uptake in the peritoneal macrophages (Dadashzadeh et al., 2010). However, the negative–negative repulsion between the negatively charged liposomes and the peritoneal surface resulted in the escape of the negative vesicles from the peritoneum. A significant uptake of the negative liposomes by the peritoneal macrophages could be the second reason for the low peritoneal level of DSPG (Sadzuka et al., 2000). An interesting finding was that the inclusion of the DSPG liposomes in the hydrogel greatly increased their peritoneal level. In

38

A. Alinaghi et al. / International Journal of Pharmaceutics 459 (2014) 30–39

spite of lowest AUC of the DSPG dispersion in the peritoneum, due to the repulsive negative–negative interaction, the highest AUC was seen in the liposomal hydrogel. The presence of anionic groups may imitate the glycerophosphate role, and thus, raise the ionic strength and enhance the electrostatic attraction between the ammonium groups of the chitosan and phosphate groups. The displacement of the glycerophosphate by negatively charged compounds such as albumin, and the binding to chitosan via the ammonium groups, was reported previously (Ruel-Gariépy et al., 2000). Further evidence was provided by the changes in the 1740 cm−1 peak of the liposome spectra (Fig. 2). The lower frequency values suggested that hydrogen banding may occur between the C O groups of the phospholipids, and OH or NH2 groups in the chitosanglycerophosphate hydrogel (Hasanovic et al., 2010). Similarly, the 3000–3600 cm−1 region, which characterizes the OH and NH bands, showed negative shifting for the DSPC and DSPG. However, the higher frequency values for the DOTAP liposomal hydrogel indicated no evidence of band formation. The presence of the DSPG shifted this band to the lower wave number that indicated a stronger interaction. This is in line with the FTIR spectra for the chitosan glycerophosphate, in that the 3000–3600 regions were more pronounced when compared to each component separately (Douglas et al., 2013). The DSPC liposomal hydrogel augmented the peritoneal retention, and there were no significant changes for the DOTAP. The positively charged liposomes leaving the hydrogel rapidly may be because of the polycationic characteristic (RuelGariépy et al., 2000) of the chitosan, which resulted in the escape of positive vesicles from the hydrogel. It should be kept in mind that the hydrogel could protect the liposomes, regardless of the lipid compositions and charges, as a barrier to avoid the vesicles being uptaken by the peritoneal macrophages; therefore the peritoneal retention has been increased. In the blood, the neutral liposomal dispersion and hydrogel showed the highest AUC values. The positively or negatively charged liposomes were cleared more rapidly than the neutral, classical liposomes (Boerman et al., 2000). Although, the DOTAP showed the lowest AUC in the blood, so that it could not increase the level of radioactivity significantly (p < 0.342), the integration with the hydrogel resulted in an enhancement ratio in the blood of 6.2. It seems that when the liposomes were injected in a dispersion form, they left the cavity very rapidly, and cleared from the blood due to the charged surface. However, when they were injected as a hydrogel, they were slowly introduced into the blood, and a higher AUC value was observed. After the IP injection of the charged liposomal dispersion, the accumulation in the GI tract, heart, and kidneys was not remarkable. A significant accumulation in the lungs was observed with the positively charged liposomes. Considering that blood cells have a negatively charged surface, one possibility was that an aggregate formed via the electrostatic interaction of the liposomes and blood cells, which were trapped in the lung capillaries (Ishiwata et al., 2000). The AUC values in the abdominal organs, such as the intestines, stomach, and kidneys, for the liposomal hydrogel formulations, increased when compared to the liposomal dispersion. One should bear in mind that due to the higher residence time of the liposomal hydrogel, 99m Tc-HMPAO was released from the liposomes, and since it has a higher affinity for accumulation in the GI tract (or elimination via the kidneys), the AUC values in these organs increased.

5. Conclusion Liposomes were integrated into the hydrogel, with the aim of controlling the release of the particulate system. Additionally, using this mixed delivery system benefits local drug retention, and a

sustained release of the drug can be achieved. Applying different liposomes as embedded particles into a polymeric-based system as a reservoir allows us to design the most effective formulation. In this study, different types of liposomes were incorporated into chitosan-glycerophosphate in situ, forming hydrogel to control the release of intact vesicles. DSPC, with a high transition temperature that made the bilayer stronger than DPPC or DMPC, presented in the blood and was retained in the peritoneum after IP injection in mice, for a longer period of time. The disposition of charged liposomes has been affected considerably due to the interaction between chitosan-glycerophosphate and the head group of the phospholipids. Negatively charged liposomes (DSPG), left the peritoneum rapidly when injected in the dispersion form, while using hydrogel could prolong the retention time in the cavity at the highest level. With respect to specific applications, it is possible to design new mixed drug delivery systems based on the characteristics of either the particulate or polymeric depot system. Acknowledgment This study was part of PhD thesis of Azadeh Alinaghi supported by Tehran University of Medical Sciences (grant No. 8560). References Alinaghi, A., Rouini, M.R., Johari Daha, F., Moghimi, H.R., 2013. Hydrogel-embeded vesicles, as a novel approach for prolonged release and delivery of liposomes, in vitro and in vivo. J. Liposome Res. 23, 235–243. Banerjee, R., 2001. Liposomes applications in medicine. J. Biomater. Appl. 16, 3–21. Boerman, O.C., Laverman, P., Oyen, W.J.G., Corstens, F.H.M., Storm, G., 2000. Radiolabeled liposomes for scintigraphic imaging. Prog. Lipid Res. 39, 461–475. Budai, L., Hajdu, M., Budai, M., Grof, P., Beni, S., Noszal, B., Klebovich, I., Antal, I., 2007. Gels and liposomes in optimized ocular drug delivery: studies on ciprofloxacin formulations. Int. J. Pharm. 343, 34–40. Charrois, G.J.R., Allen, T.M., 2004. Drug release rate influences the pharmacokinetics, biodistribution, therapeutic activity, and toxicity of pegylated liposomal doxorubicin formulations in murine breast cancer. Biochim. Biophys. Acta 1663, 167–177. Chen, L.-C., Chang, C.-H., Yu, C.-Y., Chang, Y.-J., Hsu, W.-C., Ho, C.-L., Yeh, C.-H., Luo, T.-Y., Lee, T.-W., Ting, G., 2007. Biodistribution, pharmacokinetics and imaging of 188 Re-BMEDA-labeled pegylated liposomes after intraperitoneal injection in a C26 colon carcinoma ascites mouse model. Nucl. Med. Biol. 34, 415–423. Chen, C., Han, D., Cai, C., Tang, X., 2010. An overview of liposome lyophilization and its future potential. J. Control. Release 142, 299–311. Chung, Y.M., Yang, M.C., Tsai, W.J., 2006. A fibrin encapsulated liposomes-in-chitosan matrix (FLCM) for delivering water-drugs: influences of the surface properties of liposomes and the crosslinked fibrin network. Int. J. Pharm. 311, 122–129. Dadashzadeh, S., Mirahmadi, N., Babaei, M.H., Vali, A.M., 2010. Peritoneal retention of liposomes: effects of lipid composition, PEG coating and liposome charge. J. Control. Release 148, 177–186. Douglas, T., Skwarczynska, A., Modrzejewska, Z., Balcaen, L., Schaubroek, D., Lycke, S., Vanhaecke, F., Vandenabeele, P., Dubruel, P., Jansen, J.A., Leeuwenburgh, S.C.G., 2013. Acceleration of gelation and promotion of mineralization of chitosan hydrogels by alkaline phosphatase. Int. J. Biol. Macromol. 56, 122–132. Drummond, D.C., Meyer, O., Hong, K., Kirpotin, D.B., Papahadjopoulos, D., 1999. Optimizing liposomes for delivery of chemotherapeutic agents to solid tumors. Pharmacol. Rev. 51, 691–743. Filipovic-grcic, J., Skalko-basnet, N., Jalsenjak, I., 2001. Mucoadhesive chitosancoated liposomes: characteristics and stability. J. Microencapsul. 18, 3–12. Flessner, M.F., Dedrick, R.L., Schultz, J.S., 1984. A distributed model of peritonealplasma transport: theoretical considerations. Am. J. Physiol. 246, R597–R607. Flessner, M.F., 2007. Intraperitoneal drug therapy: physical and biological principles. Cancer Treat. Res. 134, 131–152. Guo, J., Ping, Q., Jiang, G., Huang, L., Tong, Y., 2003. Chitosan-coated liposomes: characterization and interaction with leuprolide. Int. J. Pharm. 260, 167–173. Hara, M., Miyake, J., 2001. Calcium alginate gel-entrapped liposomes. Mater. Sci. Eng. C 17, 101–105. Hasanovic, A., Hollick, C., Fischinger, K., Valenta, C., 2010. Improvement in physicochemical parameters of DPPC liposomes and increase in skin permeation of aciclovir and minoxidil by the addition of cationic polymers. Eur. J. Pharm. Biopharm. 75, 148–153. Howell, S.B., 2008. Pharmacologic principles of intraperitoneal chemotherapy for the treatment of ovarian cancer. Int. J. Gynecol. Cancer 18 (Suppl. 1), 20–25. Hurler, J., Berg, O.A., Skar, M., 2012. Improved burns therapy: liposome in hydrogel delivery system for mupirocin. J. Pharm. Sci. 101, 3906–3915. Immordino, M.L., Dosio, F., cattel, L., 2006. Stealth liposomes: review of the basic science, rationale, and clinical applications, existing and potential. Int. J. Nanomed. 1, 297–315.

A. Alinaghi et al. / International Journal of Pharmaceutics 459 (2014) 30–39 Ionov, M., Gardikis, K., Wróbel, D., Hatziantoniou, S., Mourelatou, H., Majoral, J.P., Klajnert, B., Bryszewska, M., Demetzos, C., 2011. Interaction of cationic phosphorus dendrimers (CPD) with charged and neutral lipid membranes. Colloids Surf. B 82, 8–12. Ishiwata, H., Suzuki, N., Ando, S., Kikuchi, T., Kitagawa, T., 2000. Characteristics and biodistribution of cationic liposomes and their DNA complexes. J. Control. Release 69, 139–148. Jiang, C.H., Gamarnik, A., Tripp, C.P., 2005. Identification of lipid aggregate structures on TiO2 surface using headgroup IR bands. J. Phys. Chem. B 109, 4539–4544. Li, L.C., Tian, Y., 2006. Zeta potential. In: Swarbrick, J. (Ed.), Encyclopedia of Pharmaceutical Technology. , 3rd edition. Informa Healthcare, New York, pp. 4117–4128. Mirahmadi, N., Babaei, M.H., Vali, A.M., Johari Daha, F., Kobarfard, F., Dadashzadeh, S., 2008. 99m Tc-HMPAO labeled liposomes: an investigation into the effects of some formulation factors on labeling efficiency and in-vitro stability. Nucl. Med. Biol. 35, 387–392. Mirahmadi, N., Babaei, M.H., Vali, A.M., Dadashzadeh, S., 2010. Effect of liposome size on peritoneal retention and organ distribution after intraperitoneal injection in mice. Int. J. Pharm. 383, 7–13. Mufamadi, M.S., Pillay, V., Choonara, Y.E., Du Toit, L.C., Modi, G., Naidoo, D., Ndesendo, V.M.K., 2011. A review on composite liposomal technologies for specialized drug delivery. J. Drug Deliv., 1–19. Mulik, R., Kulkarni, V., Murthy, R.S.R., 2009. Chitosan-based thermosensitive hydrogel containing liposomes for sustained delivery of cytarabine. Drug Dev. Ind. Pharm. 35, 49–56. Neirinckx, R.D., Canning, L.R., Piper, I.M., Nowotnik, D.P., Pickett, R.D., Holmes, R.A., Volkert, W.A., Forster, A.M., Weisner, P.S., Marriott, J.A., Chaplin, S.B., 1987.

39

Technetium-99m d, l-HMPAO: a new radiopharmaceutical for SPECT imaging of regional cerebral blood perfusion. J. Nucl. Med. 28, 191–202. Nie, S., Hsiao, W.L., Pan, W., Yang, Z., 2011. Thermoreversible Pluronic® F127-based hydrogel containing liposomes for the controlled delivery of paclitaxel: in vitro drug release, cell cytotoxicity, and uptake studies. Int. J. Nanomed. 6, 151–166. Phillips, W.T., Rudolph, A.S., Goins, B., Timmons, J.H., Klipper, R., Blumhardt, R., 1992. A simple method for producing a Technetium-99m-labeled liposome which is stable in vivo. Int. J. Radiat. Appl. Instrum. B: Nucl. Med. Biol. 19, 539–547. Ruel-Gariépy, E., Chenite, A., Chaput, C., Guirguis, S., Lerox, J.-C., 2000. Characterization of thermosensitive chitosan gels for the sustained delivery of drugs. Int. J. Pharm. 203, 89–98. Sadzuka, Y., Hirota, S., Sonobe, T., 2000. Intraperitoneal administration of doxorubicin encapsulating liposomes against peritoneal dissemination. Toxicol. Lett. 116, 51–59. Stenekes, R.J.H., Loebis, A.E., Fernandes, C.M., Crommelin, D.J.A., Hennink, W.E., 2001. Degradable dextran microspheres for the controlled release of liposomes. Int. J. Pharm. 214, 17–20. Stewart, J.C., 1980. Colorimetric determination of phospholipids with ammonium ferrothiocyanate. Anal. Biochem. 104, 10–14. Sultana, J., Jain, R., Agil, M., Ali, A., 2006. Review of ocular drug delivery. Curr. Drug Deliv. 3, 207–217. Tsai, M., Lu, Z., Wang, J., Yeh, T., Wientjes, M.G., Au, J.L.-S., 2007. Effects of carrier on disposition and antitumor activity of interperitoneal paclitaxel. Pharm. Res. 24, 1691–1701. Zahedi, P., De Souza, R., Piquette-Miller, M., Allen, C., 2009. Chitosan-phospholipid blend for sustained and localized delivery of docetaxel to the peritoneal cavity. Int. J. Pharm. 377, 76–84.