Comparison of liposomal drug formulations for transdermal iontophoretic drug delivery

European Journal of Pharmaceutical Sciences 106 (2017) 294–301

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Comparison of liposomal drug formulations for transdermal iontophoretic drug delivery

MARK

K. Malinovskaja-Gomeza,⁎, S. Espuelasb, M.J. Garridob, J. Hirvonena, T. Laaksonenc,d a

Division of Pharmaceutical Chemistry and Technology, Faculty of Pharmacy, University of Helsinki, P.O. Box 56, FIN-00014 Helsinki, Finland Department of Pharmacy and Pharmaceutical Technology, School of Pharmacy, University of Navarra, 31080 Pamplona, Spain c Division of Pharmaceutical Biosciences, Centre for Drug Research, Faculty of Pharmacy, University of Helsinki, P.O. Box 56, FIN-00014 Helsinki, Finland d Department of Chemistry and Bioengineering, Tampere University of Technology, P.O. Box 541, FI-33101 Tampere, Finland b

A R T I C L E I N F O

A B S T R A C T

Keywords: Liposome NSAID delivery Iontophoresis Transdermal drug delivery Diclofenac sodium Skin permeation

This study was aimed to evaluate the in vitro transdermal direct/pulsed current iontophoretic delivery of an amphiphilic model compound from various lipid vesicle-encapsulated formulations compared to free-drug formulation. Conventional, pegylated, ultradeformable liposomes (transfersomes) and ethosomes loaded with a negatively charged drug diclofenac sodium (DS) were prepared and characterized. All the liposomes possessed an average size of ≈100–150 nm and negative zeta potential. No changes in colloidal stability were detected after 8 h incubation of any vesicle formulation under constant or pulsed iontophoretic current. DS was released from all the liposome formulations with a similar, limited rate (≈50% in 24 h), leading therefore to significantly lower transdermal fluxes across full-thickness porcine skin compared to the respective free drug formulation. From the tested lipid vesicle formulations, the transfersomes resulted in the highest passive flux and the ethosomes in the highest iontophoretic flux under direct constant current treatment. Higher negative surface charge of the vesicle led to better transport efficiency due to the higher mobility of the drug carrier under electric field. Pulsed current iontophoresis had no advantage over constant current treatment in combination with any type of lipid vesicular nanocarriers, in contrast to what has been described earlier with drug-loaded polymeric nanocarriers.

1. Introduction Transdermal delivery of drugs across the skin to the systemic circulation provides a convenient administration route for a variety of clinical indications (Pastore et al., 2015). In addition to avoiding the hepatic first-pass effect and chemical degradation of drugs in the gastrointestinal tract, patient compliance can be improved by reducing the frequency of dosing due to the continuous drug input. Despite of being an attractive alternative to oral and parenteral administration, however, transdermal and topical delivery of therapeutics has been clinically realized only for a handful of drugs, owing to the formidable barrier properties of stratum corneum, the outermost layer of skin. Therefore, only a limited number of molecules with appropriate balance of hydro −/lipophilicity, small size, no charge, and relatively high potency are able to pass this layer passively in therapeutic amounts (Kalia et al., 2004). In order the expand the range of molecules being able to overcome such resistance, strategies have been developed to improve the transport across or into the skin by enhancing the

⁎

permeability properties of the stratum corneum or providing a driving force acting directly on the drug (Herwadkar and Banga, 2012; Parhi et al., 2012). One such technique, iontophoresis, involves an application of mild electric current to deliver ionized or polar molecules across biological membranes (Hirvonen, 2005). The drug dose delivered by iontophoresis is directly proportional to the amount of charge passed through the skin and can be therefore controlled by the electric input of the iontophoretic system: current density, type of the current and the application time of the iontophoretic treatment. The efficiency of drug transport by iontophoresis can therefore be described by a transport number (td) that reflects the proportion of the current carried by the drug, as compared to other migrating species in formulation, and is determined by its mobility (μd), charge (zd) and concentration (cd) (Phipps and Gyory, 1992):

td =

μd ⋅z d⋅cd ∑ μ i ⋅ z i ⋅c i i

Corresponding author. E-mail addresses: kristina.malinovskaja@helsinki.fi (K. Malinovskaja-Gomez), [email protected] (S. Espuelas), [email protected] (M.J. Garrido), jouni.hirvonen@helsinki.fi (J. Hirvonen), timo.laaksonen@helsinki.fi (T. Laaksonen). http://dx.doi.org/10.1016/j.ejps.2017.06.025 Received 7 March 2017; Received in revised form 12 June 2017; Accepted 14 June 2017 Available online 15 June 2017 0928-0987/ © 2017 Elsevier B.V. All rights reserved.

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carrier, or (Caddeo et al., 2013) liposome type (conventional vs. special vesicular systems) on the permeation of loaded model drug under iontophoretic delivery, and (Cagdas et al., 2011) to determine the suitable iontophoretic current (constant vs. pulsed) type to be combined with the DS loaded vesicles.

Iontophoresis is usually carried out by a continuous direct current (DC) that is typically considered to be the most efficient current type in transdermal delivery. However, it has been suggested that the longterm use of DC may have the problem of creating a polarizing current that will decrease the efficiency of DC applied (Lawler et al., 1960). This could be overcome by a current delivered in a periodic manner (pulsed current; PC) that allows skin to depolarize and return to its initial state before the onset of next pulse (Zakzewski et al., 1992). PC is also considered to be less damaging to skin and cause less patient discomfort (Clemessy et al., 1994). Furthermore, compared to conventional DC, PC has been more effective in promoting transdermal transport of large drug-like peptides or proteins (Chien et al., 1989; Knoblauch and Moll, 1993; Malinovskaja et al., 2014; Raiman et al., 2004; Singh et al., 1998), or when the compound of interest has been loaded into polymeric nanoparticles prior administration (Malinovskaja-Gomez et al., 2016). Alternatively, in the search of improved dermal or transdermal delivery, attempts have been made to design new nanosized carrier systems, such as liposomes, micelles, nanoparticles, nanoemulsions and dendrimers, in order to ensure adequate penetration into or across the skin (Cevc and Vierl, 2010). Among those, liposomal carriers have shown to be promising drug delivery systems to transport therapeutics mainly to the different layers of skin, and by special vesicular systems (e.g. niosomes, ethosomes, transfersomes etc.) also for systemic drug delivery purposes (Pierre and Dos Santos Miranda Costa, 2011). Potential advantages of liposome use include enhanced drug delivery, solubilization of poorly soluble drugs, drug protection against proteolytic degradation, local skin depot for sustained release, reduction of side-effects and incompatibilities, or formation of rate-limiting barrier for systemic absorption (Weiner et al., 1994). There are many reports on the separate use of liposomes and iontophoresis for skin penetration enhancement, however, the combined use of both approaches has gained little attention (Essa et al., 2002a, 2002b, 2004; Fang et al., 1999; Han et al., 2004; Kajimoto et al., 2011; Kigasawa et al., 2012; Vutla et al., 1996). Combining transdermal iontophoresis with liposome-encapsulated formulations could offer some additional benefits, including improvement of drug delivery by liposome membrane/surface charge modifications and more predictable and controlled drug transport, resulting in drug fluxes less dependent on skin variables. The model drug used in this study was diclofenac sodium (DS), the most widely prescribed non-steroidal anti-inflammatory drug (NSAID) worldwide for the management of acute conditions of inflammation and pain, musculoskeletal disorders, arthritis and dysmenorrhea (Altman et al., 2015). Although widely used by oral administration, alternative delivery approaches would desirable due to the gastrointestinal side effects, such as gastric ulcers and bleeding, extensive first pass metabolism, and short biological half-life. Therefore, topical DS preparations have been developed with the aim of treating local pain and inflammation while limiting systemic exposure and potentially minimizing the risk of side effects associated with the treatment of oral NSAIDs. The physicochemical parameters (small molecular weight, lipophilic nature of the DS form while its salts are water soluble at neutral pH) makes it an excellent candidate drug to be used for transdermal delivery. Although effective passive delivery of DS across the skin has been demonstrated, the iontophoretic administration of this drug should be investigated as means to improve both the rate and extent of drug delivery, as well to decrease the variability of transdermal fluxes. The objective of this study was to test drug delivery systems that combine drug-loaded lipid vesicles and iontophoresis for the controlled transdermal delivery of small molecular weight hydrophilic model compound. In more detail, we aimed: (Altman et al., 2015) to develop a range of different DS-loaded liposome formulations suitable for transdermal iontophoretic administration, with regards to drug loading, colloidal properties, electrochemical stability and drug release kinetics, (Bahia et al., 2010) to study the effect of the surface charge of liposomal

2. Materials and methods 2.1. Chemicals Soya phosphatidylcholine (Emulmetik 930; phosphatidylcholine content 95%) was obtained from Lucas Meyer Cosmetics (Champlan, France), 1,2-distearoyl-sn-glycero-3-phosphoethanolamine-N-[methoxy (polyethylene glycol)-2000] ammonium salt from Avanti Polar Lipids (Alabaster, AL, US), and cholesterol, Tween-80 and DS from Sigma Chemical Co. (St. Louis, MO, USA). All other chemicals were at least of analytical grade. Deionized water (≥ 18.2 MΩ/cm of resistance; Millipore, Molsheim, France) was used to prepare all the solutions.

2.2. Preparation of the liposomes The conventional liposomes, the pegylated liposomes and the transfersomes were prepared by the classic cast film method. Briefly, phosphatidylcholine (PC), cholesterol (Chol), 1,2-distearoyl-sn-glycero3-phosphoethanolamine-N-methoxy-polyethylene glycol-2000 ammonium salt (DSPE-PEG2000) and Tween-80 were dissolved in absolute ethanol in a clean, dry, round-bottom flask. The organic solvent was removed by evaporation for 20 min at 45 °C to obtain homogenous thin lipid film on the inner surface of the flask. The deposited film was then hydrated with 2.5 mg/ml DS solution in 10 mM Hepes buffered saline pH = 7.4 for 1 h at 45 °C. The resulting vesicles were extruded 11 times through a 100 nm polycarbonate membrane (Avanti Polar Lipids, Alabaster, AL, US) at 45 °C. In order to prepare the ethosomes, the lipids and DS were dissolved in absolute ethanol. This mixture was heated to 30 °C ± 1 °C in a water bath. Buffer (10 mM Hepes buffered saline pH = 7.4), also heated to 30 °C ± 1 °C, was added slowly as a fine stream to lipid mixture with constant stirring at 700 rpm in a closed vessel. Mixing was continued for an additional 5 min, while maintaining the system at 30 °C ± 1 °C. The size of the vesicles was reduced by sonication for 15 s at 15 W, followed by extrusion for 11 times through the 100 nm polycarbonate membrane at ambient temperature.

2.3. Characterization of the liposomes The colloidal characteristics (hydrodynamic diameter, polydispersity index, zeta potential) of the DS loaded liposomes were determined by Malvern Zetasizer Nano (Malvern Instruments, Malvern, UK). The liposome-encapsulated DS was separated from the unentrapped drug by ultracentrifugation at 30000 rpm for 6 h at 4 °C. The liposomes were lysed with absolute ethanol and the released DS was quantified by HPLC. The percent of encapsulation efficiency (EE%) was then calculated according to the following equation:

EE% =

amount of drug in liposomes × 100 total amount of drug

(2)

The concentration of phospholipids in each formulation was quantified using the phosphate assay method (Rouser et al., 1970). The stability of the liposomes under 100% constant direct current (DC) and 75% on/25% off pulsed current (PC; cathodal, current density 0.5 mA/ cm2, frequency of pulsing 500 Hz) profiles for 8 h was evaluated at 37 °C in small glass vials as a change in hydrodynamic diameter and polydispersity index (PDI). 295

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2.4. Drug release and in vitro permeation across full-thickness porcine skin

3. Results and discussion

The drug release and the in vitro permeation across full-thickness porcine skin with or without iontophoretic current were studied in occlusive (closed) conditions in static Franz diffusion cells. The following experimental parameters were applied: cell temperature: 37 °C; the area of exposed epidermis: 0.785 cm2; donor formulation: 0.5 ml of DS solution (2 mg/ml) or drug loaded liposomes with the equivalent amount of drug and receiver formulation: 5 ml of 10 mM pH = 7.4 HEPES-buffer containing 154 mM NaCl. Both in drug release and permeation the liposome formulations were applied after removal of the non-encapsulated drug and then resuspended in the same buffer. The phospholipid contents of the formulations applied in permeation experiments were 54.00 mM, 52.24 mM, 57.63 mM and 41.24 mM for F1, F2, F3 and F4 formulations, respectively; and 10 times lower for the same formulations used in release studies. As a negatively charged drug at pH = 7.4, DS was iontophoresed from the cathodic compartment. In release experiments, a synthetic membrane (regenerated cellulose, MWCO = 6–8 kDa with ≈2 nm pore size; Spectra/Por, Spectrum Laboratories, US) was used instead of the skin to separate the different compartments of the diffusion cells. In iontophoretic permeation studies either 100% constant direct current (DC) or 75% on/25% off pulsed current (PC) was applied for 8 h, whereafter passive flux was followed up to 24 h (current density 0.5 mA/cm2; pulsing frequency 500 Hz). The skin preparation, diffusion cell set-up and iontophoretic apparatus utilized has been described in detail in our earlier study (Malinovskaja-Gomez et al., 2016).

3.1. Properties of the liposomes In this study, our goal was to study drug delivery from lipid vesicle formulations combined with transdermal iontophoresis. Therefore, we aimed at producing liposomes possessing a surface charge of the same sign as the drug loaded inside the vesicles as then the transport of the drug across the skin barrier would benefit from the iontophoretic current treatment in double. It is generally agreed that liposomes do not penetrate the skin but rather remain in the upper layers of stratum corneum (Elsayed et al., 2007). According to our hypothesis, firstly, the electric current would push the liposomes into hair follicles, that has been recognized as one the possible routes in transdermal absorption of drugs (Lauer et al., 1995). Thereafter, iontophoresis would pump the drug released from the depot in the skin or on the surface of the skin further into the bloodstream. In order to optimize this delivery technique, the effect of various liposome formulations had to be studied in a systematic way. The compositions of our final formulations used in the transdermal permeation experiments were based on the findings from literature with minor modifications. In order to study the effect of surface charge of vesicles on DS transport under iontophoresis, a pegylating agent (2.5% of molarity of total phospholipids) was used to decrease the highly negative surface charge of the conventional liposomes. DSPE-PEG2000 forms a shielding layer on the outside of the membrane of the liposome, which would cover the negative charge of the other constituents − PC and Chol (Immordino et al., 2006). Tween-80 (in 15% w/w concentration) was chosen as edge activator for the production of the transfersomes as it has been demonstrated to provide the maximum deformability to the DS vesicle membrane and transdermal flux compared to bile salts and Spans (El Zaafarany et al., 2010). For the preparation of the ethosomes, 20% of ethanol v/v was incorporated into the vesicles as this ratio has proven earlier to provide the highest transdermal fluxes of DS from ethosomes, phenomenon claimed to be related to viscous state of that dispersion that remained best adhered to skin surface (Jain et al., 2015). The average hydrodynamic diameters of the prepared liposomes ranged from 116.5 to 147.0 nm, with the conventional and the pegylated liposomes possessing slightly higher sizes than the transfersomes and the ethosomes (Table 1). These results were expected, considering the differences in composition and manual extrusion, and were not believed to lead to significant changes in transdermal drug fluxes. Still, there is lot of controversy among reports on the influence of liposome particle size on the drug transport across the skin. While some studies have not found higher disposition of drug either in lower skin strata or receiver compartments irrespective of when the drug was applied in small or large vesicles (Du Plessis et al., 1994; Kajimoto et al., 2011), others claim the opposite (Schramlova et al., 1997; Verma et al., 2003).

2.5. Drug analysis The DS concentrations were analyzed by HPLC using the following conditions: column Supelco Discovery® C18, 5 μm; 150 ∗ 4.6 mm; mobile phase: phosphate buffer pH = 3: methanol (30:70); flow rate: 1 ml/min; wavelength: 284 nm; injection volume: 50 μl; column temperature: 25 °C and retention time: 3.3 ± 0.2 min. The linear calibration curve (r2 = 0.9999) was established in the range of 0.1–100 μg/ml of drug concentration. LOQ = 0.1 μg/ml and %RSD was 0.13% and 0.18% for retention times and peak areas, respectively.

2.6. Data analysis The linear part of the permeation curve was used for calculating the steady-state flux values (nmol/h ∗ cm2). All the results were expressed as means ± SD and unpaired t-test or one-way ANOVA with Bonferroni's multiple comparison test was used for statistical analysis, when necessary.

Table 1 Composition, particle size, polydispersity index (PDI), zeta potential and mean encapsulation efficiency (EE%) of prepared vesicles. Formulation and code

Composition

Size (nm)

PDI

Charge (mV)

Mean EE %

Conventional liposomes (F1)

90 mM PC 30 mM Chol 90 mM PC 30 mM Chol 2.5% (mol/mol%) DSPE-PEG2000

147.0 ± 1.5

0.085 ± 0.013

− 34.3 ± 1.7

67.5

144.2 ± 0.1

0.079 ± 0.029

− 27.7 ± 1.0

60.4

117.7 ± 0.5

0.073 ± 0.028

− 29.8 ± 1.6

79.9

116.5 ± 0.6

0.104 ± 0.005

− 26.7 ± 1.9

49.6

Pegylated liposomes (F2)

Transfersomes (F3)

Ethosomes (F4)

120 mM PC 15% (w/w%) Tween-80 90 mM PC 30 mM Chol 20% (v/v%) ETOH

PC– soya phosphatidylcholine (Emulmetik 930); Chol – cholesterol; DSPE-PEG2000–1,2-distearoyl-sn-glycero-3-phosphoethanolamine-N-methoxy-polyethylene glycol-2000 ammonium salt; DS – diclofenac sodium.

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et al., 1999). The antimicrobial drug enoxacin-loaded vesicles showed no changes in drug encapsulation under constant current treatment for 6 h and the size of the current-treated liposomes was generally even smaller than of the untreated liposomes after incubation in buffer. Although pore widening of lipid bilayers under electric field has been associated with the breakdown of liposomes and the leakage of the loaded drug, the magnitude of current used in typical transdermal iontophoretic treatment regimens is too small (maximum current density is typically 0.5 mA per cm2 of skin) to cause rupture of the vesicles. Furthermore, under electric current field, the fusion of charged vesicles is inhibited, so the current can even contribute to better colloidal stability of the formulations.

The particle size distribution of all our liposomes was monomodal with a PDI value of ≈ 0.1, indicating a narrow size distribution and an excellent dispersion homogeneity. The zeta potentials were measured with the purified liposomes, after being dispersed in 1 mM KCl solution. All the liposomes were negatively charged with the conventional liposomes possessing the highest negative zeta potential value (−34.5 mV). The addition of the pegylated phospholipid (2.5% of phospholipid molarity) to the formulation decreased the charge (lowered the negative zeta potential value) to − 27.7 mV. Also the transfersomes and the ethosomes had slightly lower negative surface charges compared to the conventional liposomes. Encapsulation efficiency shows the percentage of the initial drug incorporated into liposomes. Results showed that all the vesicles were able to encapsulate DS with good yields (EE% values ranged from 46.6% to 79.9%), with the highest EE% reached in the transfersomes. As DS possesses amphiphilic properties, it can be encapsulated in both, bilayer and aqueous compartment of the liposomes, and the overall amount of loaded DS therefore corresponds to the fractions of the drug incorporated in both phases. Amphiphilic compounds have been shown to interact with the phospholipid bilayer of vesicles, leading to modifications in membrane properties, high drug loading or even phase transition to micellar solution (Moreno et al., 2009; Schütze and MüllerGoymann, 1998). For DS, it has been proposed, that carboxyl function of diclofenac anion interacts with the positively charged ammonium group of choline changing the electronic density of phosphorus atom (Lopes et al., 2004). The high DS encapsulation efficiencies obtained in our study could be explained by these interactions.

3.3. Release of DS from the liposomes The drug formulation to be combined with transdermal iontophoresis should ideally avoid immediate/burst drug release and deliver the drug to the permeation site in a gradual yet continuous manner. In this way, more predictable and controlled drug transport could be achieved, resulting in fluxes less dependent on skin variables. In this study, the release kinetics of DS was studied in a similar setting as the in vitro transdermal permeation − at 37 °C in static Franz diffusion cells, only using a synthetic dialysis membrane to separate the donor compartment from the receiver compartment. Fig. 2 illustrates the release profiles of DS from the different liposome formulations. As a control, the same amount of drug dissolved in the buffer was utilized. From all the lipid vesicle formulations DS was released with similar rate, resulting in ≈ 50% of the encapsulated drug released in 24 h. In general, the most important factors affecting drug release from liposomes are the physicochemical properties of the liposome membrane and the therapeutic agent, and the vesicle size (Lindner and Hossann, 2010). Overall, the strong binding/interactions of DS to/with phospholipid bilayer explains the slow drug release from liposomes as the drug partitioning from the lipid phase becomes the determining factor over release kinetics (Lopes et al., 2004; Moreno et al., 2009). The incorporation of cholesterol leads typically to the formation of more ordered and rigid lipid membranes (Cagdas et al., 2011), whereas surfactants (edge activators in transfersomes) and ethanol result in leakier vesicle membranes, which makes the drug release easier (El Zaafarany et al., 2010; Rakesh and Anoop, 2012). All the studied formulations possessed similar vesicle sizes and drug-to-lipid ratios (data not shown). Thus, the effect of these composition parameters on drug release was not evident in this study. In summary, all the formulations demonstrated suitable release kinetics to be used in transdermal delivery and any difference in DS transport efficiency across the skin from our liposome formulations could only be attributed to the differences in liposome-skin-iontophoretic current interactions, not in drug release

3.2. Electrochemical stability of the liposomes Prior to the permeation experiments, the colloidal stability of the vesicles was evaluated under direct constant and pulsed current modes. The results presented as a change in hydrodynamic diameter and PDI are shown in Fig. 1. No changes in hydrodynamic diameters were detected after 8 h incubation in any vesicle formulation under constant or pulsed current. The PDIs of the formulations F1, F2 and F3 had increased slightly, but these changes did not reach statistical significance (p > 0.05 when compared with unpaired t-test before and after current treatment). As all the tested liposomes were stable under iontophoretic current, the formulations were deemed suitable for subsequent transdermal permeation studies. Only one earlier study, with regards to the combined use of liposomes and transdermal iontophoresis, has been reported on the electrochemical stability of liposomes under iontophoretic current (Fang

Fig. 1. Electrochemical stability of liposomes under iontophoretic current for 8 h. Expressed as a change in hydrodynamic diameter and PDI. Mean ± SD, n = 3.

Fig. 2. Release of diclofenac sodium (DS) from different liposome formulations (F1–F4). The total amount of DS in liposomes was 100 μg. Mean ± SD, n = 3.

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membrane (the total amount of current used in the PC treatment was 75% of the DC treatment).

rates. 3.4. Transdermal in vitro delivery of DS across porcine full-thickness skin

3.4.2. Delivery from the liposomes This is the first report on investigating the transdermal iontophoretic delivery of DS from liposomal carriers. In general, the use of DS loaded vesicles in the donor compartment led to decreased flux values and slightly longer lag-times to reach steady-state fluxes compared to DS delivery from solution. This phenomenon was caused by gradual and limited release of drug from the liposomes during the 24 h permeation experiments. The inability of lipid vesicles to enhance transdermal DS permeation compared to free-drug formulation has been also demonstrated earlier (Caddeo et al., 2013; Fathi-Azarbayjani et al., 2015). This could be explained by the hydrophilic nature of the compound that leads to poor drug partitioning from the vesicles into lipid environment of the skin. The application of the iontophoretic current enhanced the delivery of our model drug compared to passive delivery in a similar magnitude than from the solution. On Fig. 3B the permeation curves of the liposome formulations that provided the highest passive and iontophoretic flux values are presented. From all the tested lipid vesicles, the ultradeformable liposomes (F3) resulted in the highest passive flux and the ethosomes (F4) the highest iontophoretic flux under direct constant current treatment, although the analysis by one-way ANOVA test failed to demonstrate statistical difference. One of the aims of this study was also to study the effect of the surface charge of the liposomal nanocarriers on the model drug transport under iontophoretic current. For that purpose we compared the DS transport from the pegylated liposomes with reduced negative zeta potential value (F2) to the highly negatively charged conventional liposomes (F1). As expected, the highly negatively charged conventional liposomes led to more efficient iontophoretic delivery of DS, indicating that liposome charge can affect the iontophoretic treatment outcome. Also, the enhancement factors and the transport numbers of F2 were lower than of F1. At the same time, the passive drug fluxes from both of the formulations were in the same range. This indicates that the higher iontophoretic flux of F1 liposomes could be related directly to more pronounced mobility of the nanocarrier under electric current (electrorepulsion of liposome) leading to improved diffusion of the liposomes − not to changes in skin-liposome interactions. In previous reports, other techniques besides pegylation have been used to change/

3.4.1. Delivery from solution There have been several reports on transdermal iontophoretic delivery of DS either in vitro (Fang et al., 2000; Fang et al., 1999; Kasha et al., 2012; Koizumi et al., 1990) or in vivo (Clijsen et al., 2015; Crevenna et al., 2015; Garcia et al., 2016; Riecke et al., 2011; Varghese and Khar, 1996). Also combinations of iontophoresis with other chemical or physical enhancement techniques for the delivery of DS have been utilized (Kigasawa et al., 2009; Patel et al., 2015; Sugibayashi et al., 2000; Xin et al., 2012). Our permeation studies were performed at pH 7.4, where DS is negatively charged (pKa = 4.15) and can therefore be administered by cathodal iontophoresis. According to the Henderson–Hasselbalch equation, the ionization degree of an acid I% = 100 / [1 + antilog (pKa − pH)] (Florence and Attwood, 2011). At pH 7.4 practically all of DS (99.94%) is in ionized state and can therefore maximally benefit from the iontophoretic current treatment. Obviously, the natural pH gradient existing between the acidic conditions on stratum corneum (pH ≈ 5) and the physiological pH 7.4 of the dermis and the excellent buffering properties of skin could have an effect on the ionization degree and thus the iontophoretic transport efficiency of DS (Darlenski and Fluhr, 2017; Riviere and Brooks, 2009). Under iontophoretic current the total flux of a charged molecule is typically the sum of fluxes driven by electrorepulsion, electroosmosis and passive diffusion (Kalia et al., 2004). The iontophoretic flux of DS is expected to be the sum of electrorepulsion and passive diffusion as electroosmosis does not contribute to the transport of negatively charged solutes and can be neglected in our case. Results from the in vitro transdermal permeation experiments with the DS solution and the liposomes are presented in Table 2 and Fig. 3. The experiments with the DS solution demonstrated that iontophoresis significantly increased the transport of the negatively charged model drug across porcine skin compared to passive transport (e.g. EDC = 5.1). As expected, the direct constant current was more effective in enhancing the delivery of DS to receiver compartment compared to pulsed current that resulted in ≈ 25% lower transdermal flux. This can be considered typical as the iontophoretic flux of charged small molecular compounds is mostly driven by electrorepulsion mechanism that is directly related to the amount of current passed across the permeation

Table 2 The results from the permeation experiments with DS solution and different liposome formulations (F1–F4). Mean, n = 5–6.

Solution

F1

F2

F3

F4

Treatment

Jssa (nmol/h ∗ cm3)

Tnb

Passive DC PC Passive DC PC Passive DC PC Passive DC PC Passive DC PC

2.78 ± 1.31⁎⁎ 14.21 ± 7.51⁎⁎⁎ 10.45 ± 4.27⁎⁎⁎ 1.11 ± 0.43 3.98 ± 1.78 2.94 ± 0.90 1.11 ± 0.56 3.64 ± 1.12 1.78 ± 0.72 1.44 ± 0.30 3.47 ± 1.23 2.38 ± 0.64 0.71 ± 0.18 4.53 ± 1.59 2.87 ± 1.82

– 8.13 × 10− 4 7.47 × 10− 4 – 2.13 × 10− 4 2.10 × 10− 4 – 1.95 × 10− 4 1.27 × 10− 4 – 1.86 × 10− 4 1.70 × 10− 4 – 2.43 × 10− 4 2.05 × 10− 4

± 3.81 × 10–4⁎⁎⁎ ± 2.86 × 10–4⁎⁎⁎ ± 9.53 × 10− 5 ± 6.46 × 10− 5 ± 7.47 × 10− 5 ± 5.09 × 10− 5 ± 6.72 × 10− 5 ± 3.98 × 10− 5 ± 1.04 × 10− 4 ± 1.39 × 10− 4

Ec

Q24hd (μg)

% of JPC of JDCe

– 5.1 3.8 – 3.6 2.6 – 3.3 1.6 – 2.4 1.6 – 6.4 4.1

22.63 ± 9.48⁎⁎⁎ 79.68 ± 25.10⁎⁎⁎ 73.92 ± 17.89⁎⁎⁎ 6.39 ± 2.64 23.76 ± 2.90 22.96 ± 4.41 11.62 ± 3.93 25.25 ± 7.60 14.89 ± 4.36 8.48 ± 3.81 24.37 ± 8.17 18.50 ± 4.57 3.68 ± 0.95 27.42 ± 7.54 19.07 ± 7.24

– – 73.5 – – 73.9 – – 48.6 – – 68.6 – – 63.4

Jss – steady-state flux, calculated from the linear slope of the permeation curve. zFdQ dt Tn - transport number, calculated using an equation: Tn = where z is the valence of the drug, F is the Faraday's constant, dQ / dt is the linear slope of the iontophoretic phase MW ⋅ I , (in mass units per time), MW is the molecular weight and I is the total current passed through the skin. c E – enhancement factor; E = Jiontophoresis / Jpassive. d Q24 h - cumulative amount of drug permeated in 24 h. e % of JPC of JDC – percent of steady-state flux of pulsed current treatment from steady-state flux of constant current treatment, mean value. ⁎⁎ p ˂ 0.01 Jss, Tn or Q24 h from solution vs. Jss, Tn or Q24 h from liposome formulations without or with respective iontophoretic treatment. ⁎⁎⁎ p ˂ 0.001 Jss, Tn or Q24 h from solution vs. Jss, Tn or Q24 h from liposome formulations without or with respective iontophoretic treatment. a

b

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Fig. 3. The in vitro transdermal permeation of DS across porcine full-thickness skin. A. The permeation curves of DS from solution formulation passively and under direct constant (DC) or pulsed current (PC). B. The liposome formulations with best passive (F3, passive) or iontophoretic flux (F4, DC). DC – 100% constant direct current 0.5 mA/cm2; PC – 75% on:25% off pulsed current. The amount of DS in the donor compartment was 1000 μg (mean ± SD, n = 5–6).

transport across the skin in passive delivery, reaching higher fluxes than from free-drug solution or commercial DS gels (Cevc and Blume, 2001; El Zaafarany et al., 2010; Ghanbarzadeh and Arami, 2013; Jain et al., 2015). In regards to the composition of transfersomes, the choice of edge activator and its concentration can have a crucial outcome on drug transport (El Zaafarany et al., 2010). We have considered that the removal of the non-encapsulated drug prior permeation studies could have resulted in partial loss of the edge-activator Tween-80 from the vesicle membrane and change in liposome membrane properties, as described earlier (Bahia et al., 2010). Also, for transfersomes, open (non-occluded) administration has been proposed to be a prerequisite to obtain maximum permeation enhancing effect, as the driving force for penetration − transdermal hydration gradient − is eliminated by occlusive conditions (Cevc and Blume, 1992). At the same time, as the application of iontophoretic current requires aqueous media around the electrodes, all our permeation experiments were conducted under occlusion to prevent the evaporation of the buffer from the donor compartment. Interestingly, occlusion did not abolish totally the penetration enhancing effects of the transfersomes compared to the other tested lipid vesicles. Similar results were obtained when estradiol transport from transfersomes was studied under occlusion (Essa et al., 2004). This is the first study on the effect of iontophoretic current type on transdermal delivery of drug loaded liposomes. In all earlier studies, where liposomes were combined with iontophoresis, only constant direct current profile had been utilized (Essa et al., 2002a, 2002b, 2004; Fang et al., 1999; Han et al., 2004; Kajimoto et al., 2011; Kigasawa et al., 2012; Vutla et al., 1996). Typically, constant current treatment is the most efficient in enhancing transdermal delivery of charged or ionic drugs, as for most compounds the fluxes are proportional to the amount of current passed across the skin. However, the use of constant current has its limitations, especially in long-term use (see Introduction). Also, for some larger drugs (peptides) or even drug-loaded polymeric carrier systems pulsed current has been proven to be more efficient in transport enhancement (Malinovskaja-Gomez et al., 2016; Malinovskaja et al., 2014; Raiman et al., 2004). Interestingly, in this study, pulsed current treatment had no clear advantage over constant current treatment in combination with any of the liposome formulations. The treatment was especially ineffective with the liposomes with lower surface charge, such as the pegylated or the ultradeformable liposomes, as was seen from the % of JPC of JDC value (that shows the percent of the steadystate flux of pulsed current treatment from the steady-state flux of constant current treatment). This demonstrated that the lipid vesicles with less mobility under electric current were benefitting less from the iontophoretic treatment delivered in pulsatile manner. On the contrary, drug delivery from the polymeric nanospheres was more enhanced under pulsed current than under constant direct current, indicating that there must be other factors than surface charge underlying the interaction between the drug-loaded nanocarriers and current type

impart charges of/to drug loaded liposomes. Cationic lipids like stearylamine, DDAB, DODAP, DOTAP, DOSPER, spermine or protamine sulfate have been added to the formulations to induce positive charge, and negative lipids such as phospatidylserine, dicetylphosphate or DMPG, have been utilized to impart a negative charge (Fang et al., 1999; Han et al., 2004; Vutla et al., 1996). These earlier studies have demonstrated the following: firstly, for effective drug transport, the surface charge of the liposome has to be of the same sign as the drug loaded inside; secondly, the charge inducing excipient is more efficient if it is a multi-cation by nature and does not possess too big molecular size as this lowers the mobility under electric current, and thirdly, a possible competition effect between the charge inducing agent and the drug is possible, especially when the excipient is a smaller molecule and therefore more mobile under electric current. Considering the possible competitive ion effect that might have an unwanted effect on drug flux, we chose pegylation in this study for the charge effect comparison. Although pegylation was demonstrated to be a good technique in shielding the negative charge of vesicle, its potential remained somewhat limited. In general, it is agreed, that conventional liposomes have not much value as drug carriers for transdermal delivery (Pierre and Dos Santos Miranda Costa, 2011) as the deep skin penetration is inhibited and the vesicles remain in the upper layer of stratum corneum. With special techniques carrier vesicles can be produced that enable modulated delivery across the skin layers and provide systemic action. In this study we investigated the potential of two such vesicle types – ultradeformable liposomes and ethosomes – to be used in combination with iontophoresis and to enhance transdermal DS delivery. Ultradeformable liposomes (transfersomes), having an edge activity in their lipid bilayer, are described as freely moving very flexible vesicles that are able to penetrate the skin due to proposed transdermal hydration force (Cevc et al., 1998). Ethosomes are known to facilitate drug permeation by a tentative synergistic “ethanol-ethosome effect” in which ethanol disrupts the organization of stratum corneum lipid bilayer, allowing the flexible ethosomes then to penetrate and possibly permeate this layer (Touitou et al., 2000). From all the tested vesicle formulations, the transfersomes led to the highest passive flux of DS across porcine skin. Due to the longer lag-time (data not shown) to reach the steady-state flux, the total amount of DS delivered during 24 h remained still smaller than from the pegylated liposomes. Interestingly, the iontophoretic current was quite inefficient in further enhancing DS delivery from the transfersomes (EDC = 2.4). The prepared ethosomes failed to increase DS permeation passively compared to the conventional liposomes, but at the same time the iontophoretic enhancement was the highest (EDC = 6.4), also leading to the highest steady-state iontophoretic flux and transport number among all the liposome formulations. In contrast to our results, DS formulations based on transfersomes and ethosomes have been demonstrated several times to significantly improve DS 299

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(Malinovskaja-Gomez et al., 2016). It is also possible, that in combination with lipid vesicles, the pulsed current is more responsible in delivering the drug into deeper layers of skin, but not enhancing the permeation further across the skin to systemic blood circulation. 4. Conclusions This study tested a system combining iontophoresis and drug-loaded lipid vesicles for the controlled transdermal delivery of an amphiphilic small molecular weight model compound. Therefore, four different types of lipid vesicles − conventional, pegylated, transfersomes and ethosomes − loaded with diclofenac sodium were prepared and the effect of differences in composition was evaluated on drug transport efficiency across the skin. Although all the obtained liposome formulations were deemed as suitable for transdermal iontophoretic administration, regarding the colloidal properties, the stability under iontophoretic current and the release kinetics, no improvement in the drug delivery was seen from nanoencapsulating diclofenac sodium into lipid vesicles compared to free drug formulation. All the liposome formulations led to reduced transdermal fluxes and smaller drug amounts delivered, both in passive and iontophoretic delivery regimens. Iontophoretic drug transport from the liposomes was significantly affected by the composition and the charge of the lipid bilayer of the vesicles, and the iontophoretic current mode utilized. The transfersomes resulted in the highest passive flux and the ethosomes in the highest iontophoretic flux under direct constant current treatment. The lipid vesicles with a higher negative surface charge led to better transport efficiencies of the loaded model drug due to the higher mobility of the drug carriers under iontophoretic current. This is the first report to study the effect of current type on the transdermal permeation of a drug loaded into lipid vesicles. Based on our findings, pulsed current treatment has no clear advantage over constant current treatment in combination with any type of lipid vesicular nanocarriers, which is in contrast to what has been described earlier with polymeric nanocarriers. Abbreviations DDAB dimethyldiooctadecylammonium bromide DODAP 1,2- dioleoyl-3-dimethylammonium propane DOSPER (1,3-dioleoyloxy-2-(6-carboxyspemyl)) propylamide (tetraammonium tetraacetate) DOTAP 1,2-dioleoyl-3-trimethylammonium propane DC direct current DS diclofenac sodium E iontophoretic enhancement factor F Faraday constant (C/mol) HEPES 4-(2-hydroxyethyl)-1-piperazineethanesulphonic acid HPLC high performance liquid chromatography JSS steady-state flux (μg/h per cm2) MW molecular weight (g/mol) MWCO molecular weight cut off PC pulsed current PDI polydispersity index Tn transport number Z valence Acknowledgements The financial support from the Academy of Finland (grant nos. 258114 and 264988) is gratefully acknowledged. References Altman, R., Bosch, B., Brune, K., Patrignani, P., Young, C., 2015. Advances in NSAID development: evolution of diclofenac products using pharmaceutical technology. Drugs 75, 859–877.

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