Colloids and Surfaces B: Biointerfaces 123 (2014) 339–344
Contents lists available at ScienceDirect
Colloids and Surfaces B: Biointerfaces journal homepage: www.elsevier.com/locate/colsurfb
Fabrication of monodisperse liposomes-in-microgel hybrid microparticles in capillary-based microfluidic devices Eun Seon Jeong a,1 , Han Am Son b,1 , Min Kyung Kim a , Kyoung-Ho Park c,∗∗ , Sechan Kay d , Pil Seok Chae a , Jin Woong Kim a,c,∗ a
Department of Bionano Technology, Hanyang University, Ansan, Gyeonggi-do 426-791, Republic of Korea Korea Institute of Geoscience and Mineral Resources, Daejeon 305-350, Republic of Korea c Department of Applied Chemistry, Hanyang University, Ansan, Gyeonggi-do 426-791, Republic of Korea d Hwa Costech Co., Seongnam, Gyeonggi-do 463-863, Republic of Korea b
a r t i c l e
i n f o
Article history: Received 4 June 2014 Received in revised form 25 August 2014 Accepted 17 September 2014 Available online 28 September 2014 Keywords: Liposome Microgel Microfluidics Hydrogel mesh size Drug releasing
a b s t r a c t This study introduces a drop-based microfluidic approach to physically immobilize liposomes in microgel (liposomes-in-microgel) particles. For this, we generate a uniform liposomes-in-water-in-oil emulsion in a capillary-based microfluidic device. Basically, we have investigated how the flow rate and flow composition affect generation of emulsion precursor drops in micro-channels. Then, the precursor emulsion drops are solidified by photo-polymerization. From characterization of hydrogel mesh sizes, we have figured out that the mesh size of the liposomes-in-microgel particles is bigger than that of bare microgel particles, since liposomes take space in the hydrogel phase. In our further study on drug releasing, we have observed that immobilization of liposomes in the microgel particles can not only remarkably retard drug releasing, but also enables a sustained release, which stems from the enhanced matrix viscosity of the surrounding hydrogel phase. © 2014 Elsevier B.V. All rights reserved.
1. Introduction The molecular geometry of the unit amphiphiles, which is typically determined by the packing parameter (P = V/aL), can be correlated with their assembled architecture: P is the packing parameter, V and L are the effective volume and length of the amphiphile hydrophobic chain, and a is the amphiphile optimal area at the interface [1,2]. For a packing parameter of ∼1/3, a micelle structure is generated. A bilayered vesicle is made, when the packing parameter is 1/2–1. For a packing parameter of ∼1, a planar bilayer, commonly referred to as lamellar, is formed. The lipid is a sort of amphiphile molecules having both hydrophilic and hydrophobic properties. The packing parameter of lipids ranges 1/2–1, so that they form a lipid membrane with different
∗ Corresponding author at: Department of Applied Chemistry, Hanyang University, 55 Hanyangdaehak-ro, Sangnok-gu, Ansan, Gyeonggi-do 426-791, Republic of Korea. Tel.: +82 31 400 5499. ∗∗ Corresponding author. Tel.: +82 31 400 5409. E-mail addresses: [email protected] (K.-H. Park), [email protected] (J.W. Kim). 1 These authors were equally contributed to this work. http://dx.doi.org/10.1016/j.colsurfb.2014.09.039 0927-7765/© 2014 Elsevier B.V. All rights reserved.
curvatures, which leads to a structurally defined vesicle geometry after the molecular assembly. Liposomes have a bilayered vesicular structure. This structural uniqueness allows them to encapsulate the hydrophilic molecules within the interior aqueous core as well as the hydrophobic molecules within the bilayered lipid shell, respectively [2]. Hence, they are able to protect the encapsulated reactive or sensitive compounds from degradation by shielding and stabilizing against environmental and chemical stresses. Thanks to this function, liposomes are of great interest in pharmaceutical, cosmetic, food and biomedical industries [3,4]. Their size, surface properties, hydrophobicity, and functionality can be tailored by tuning the chemistry and composition of phospholipids and by altering environmental conditions, such as solvent type, ionic strength, pH, and temperature [5–7]. Most of all, the abilities to deliver active and labile molecules into cells or tissues either by vesicle fusion or control of membrane permeability have provided them with more practical applications [8–11]. Despite liposomes have such extensive applicability in academy as well as in industries, they have several critical set-backs, which are currently hampering much wider uses. The carriers made with lipid molecules readily fuse into the membrane bilayers. This makes them fade away quickly upon administrating through organs [12]. Also, liposomes have tendency toward being taken away rapidly
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by immune cells in the blood. In order to resolve these problems, a number of studies have been performed. In typical, they have tailored the periphery of liposomes to physically shield the inner part from external stresses: for instance, the structural stability of liposomes can be enhanced by co-assembly of amphiphilic graft copolymers, so that they can afford to stay in the blood much longer than non-modified ones [13]. One more critical issue that limits the application of liposomes in complex formulations is that the liposomal structure is susceptible to the presence of surfactants with low packing parameters. On adding a very small amount of these surfactants, the liposome-to-micelle transition readily occurs [14,15]. To overcome the above-mentioned drawbacks, in this study, we introduce a facile and trouble-shooting approach in which liposomes were incorporated into microgel microparticles and physically immobilized in the hydrogel network. Hydrogels are able to load and release drug molecules through their web-like molecular network. In principle, the diffusion of drug molecules is mainly ruled by the viscosity of the medium phase, which can be understood by the Stokes-Einstein equation, D = kB T/6a, where D is the diffusion coefficient, kB is the Boltzmann constant, T is the absolute temperature, is the viscosity of the medium, and a is the hydrodynamic radius of the particle. This means that the release kinetics is controllable via regulation of matrix viscosity. To provide a rationale for this, we immobilize the liposomes in uniform microgel particles. This particle system is named as liposomes-in-microgel (L-i-M). For this, we produce a uniform-sized liposomes-in-water-in-oil (L/W/O) emulsion by using a capillary-based microfluidic technique and the L/W drops containing the liposomes are solidified by photopolymerization [16,17]. Basically, the hydrogel phase is made of poly (2-methacryloyloxyethyl phosphorylcholine) (PMPC) having an excellent biocompatibility [18]. Finally, we try to demonstrate that our approach to immobilize the liposomes in the hydrogel microparticles enables a matrix-mediated controlled release. 2. Experimental 2.1. Materials 1,2-Dipalmitoyl-sn-glycero-3-phosphocholine (DPPC) was mercifully offered from Doosan Co. (Korea). 2Methacryloyloxyethyl phosphorylcholine (MPC) was kindly supplied from KCI Co. (Korea). Cetyl PEG/PPG-10/1 dimethicone (Abil EM 90, Evonik, Germany) and hexyltrimethoxysilane (TCI, Japan) were used as received. Chloroform was purchased from Deajung (Korea). Isopropanol was purchased from Samcheon (Korea). Paraffin oil, glycerin, N,N -methylenebisacrylamide (BIS), 2-hydroxy-2-methylpropiophenone (Darocure 1173), fluorescein isothiocyanate (FITC), and fluorescein isothiocyanate dextran (FITC-dextran, 70 kDa) were all purchased from Sigma–Aldrich (USA). Texas Red-1,2-dihexadecanoyl-snglycero-3-phosphoethanolamine (Texas Red-DHPE) was bought from Invitrogen (USA). All other chemicals were reagent grades and used without further purification. For all experiments, deionized double distilled water was used. 2.2. Synthesis of DPPC liposomes DPPC liposomes were prepared using the modified thin-film rehydration method [19,20]. First, DPPC was completely dissolved in chloroform and evaporated with a rotary evaporate at 52 ◦ C for 2 h to remove all the traces of the organic solvent. Then, the thin DPPC film appeared on the wall of the round flask was rehydrated with water and applied mild sonication for 2 h at 65 ◦ C in
order to come off the film into water and to form micron-sized multi-lamellar structured liposomes. The concentration of DPPC in the suspension was set to 1 wt%. To further decrease the particle size and fabricate uni-lamellar liposomes, strong probe-sonication was carried out pulsing every 1 s for 10 min with a power of 130 W at room temperature (VCX130, Sonics & materials Inc.). The size of liposomes was characterized with a dynamic light scattering (ELSZ, Otsuka electronics, Japan). The morphology of liposomes in an aqueous phase was observed with a TEM (Energy-Filtering Transmission Electron Microscope, LIPRA120, Carl Zeiss, Germany). The test samples were negatively stained with 1 wt% of uranyl acetate in the aqueous solution. Then, they were dried in air before TEM observation. 2.3. Fabrication of micro-capillary microfluidic devices To fabricate a capillary-based microfluidic devices, first, a round capillary was tapered by heating and pulling a cylindrical glass capillary (outer diameter = 1.0 mm, World Precision Instruments, USA) with a pipette puller (Model P-97, Sutter Instruments, USA). The end tip of the tapered glass capillaries was cut to the designated diameter using a microforge station (Micro Forge MF 830, Narishige, Japan). To prevent any wetting of the trimmed round capillary by the aqueous inner fluid, hydrophobic coating of the round capillary was conducted with 1 wt% of hexyltrimethoxysilane in toluene. For generation of uniform emulsions, a tapered cylindrical capillary was inserted into a square capillary (Atlantic International Technology, USA). The inner diameter of the square capillary was the same as the outer diameter of the round capillary by 1 mm. Each end of the square capillary was fit into with a needle tip and completely glued with epoxy resin. 2.4. Synthesis of L-i-M microparticles To generate emulsions, each fluid was loaded into a glass syringe (Hamilton Gastight, USA) connected with a polyethylene tube (PE5, Scientific Commodities, USA). Then, the tube was connected with the needle installed at each end of the square capillary. The dispersion fluid (DF) was an aqueous monomer solution containing the liposome suspension (0–15 wt%), MPC (15 wt%), BIS (a crosslinker, 0.5–1.5 wt%), Darocure 1173 (a photo-initiator, 0.5 wt%), and glycerin (30 wt%). The outer fluid (OF) consisted of paraffin oil and 2 wt% Abil EM 90. The flow rate of each fluid was precisely controlled with a syringe pump (Pump 11 Elite, Harvard Apparatus, USA). The emulsion drops were collected through the round capillary and observed with a bright-field microscope (Samwon, NSI-100, Korea). Then, the emulsion drops were solidified by photo-polymerization under UV 365 nm for 1 min. The paraffin oil, remnant monomers, and other additives were thoroughly removed with a large amount of isopropanol by repeated centrifugation at 4000 rpm. Finally, the particles were re-dispersed in water. The volumes of the microgel particles in a confined state, a swollen state, and a collapsed state were determined by measuring their diameters after drying, before swelling, and after swelling with a bright-field microscope. To demonstrate that the liposomes were locked in the microgel particles, we encapsulated 50 L of FITC-dextran in the core and coassembled a small amount of Texas Red-DHPE with the DPPC layers, respectively. The distribution of fluorescence probe molecules in the particles was detected with an inverted fluorescence microscope (IX-81, Olympus, Japan). 2.5. Investigation of releasing profiles Release studies were carried out by using pyrene (ex 336/em 393) as a model drug. For this, a pyrene-loaded liposomal suspension was produced using the same preparation procedure.
E.S. Jeong et al. / Colloids and Surfaces B: Biointerfaces 123 (2014) 339–344
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Fig. 1. (A) A schematic of microfluidic set-up for generation of an L/W/O emulsion. (B) A schematic of a liposome particle. (C) A TEM image of the liposome made with 1 wt% DPPC. (D) A particle size distribution of the DPPC liposome.
Undissolved pyrene in the continuous aqueous solution was filtered out with a syringe filter (800 m pores). Pyrene-loaded L-i-M particles were also prepared by following the established method. The concentration of pyrene in both suspensions was regulated with 0.02 w/v%. The suspension was sealed in a dialysis membrane bag (molecular weight cut-off ∼ 10,000, Spectrum lab, Inc., USA) that was immersed into a conical tube containing 40 mL water. Then, the conical tube was gently tumbled with a mixer (RW20D, IKA, Germany) at 42 ◦ C. 150 L of the outer aqueous phase was sampled with the time interval. After each sampling, the outer phase was replaced with the same volume of water. The concentration of pyrene released out was monitored by detecting the intensity of fluorescence with a fluorescence spectroscope (RF-5301PC, Shimadzu, Japan).
process physically fixed the liposomes in hydrogel matrix. The phospholipid used to prepare the liposomes in this study was DPPC (Fig. 1B). To obtain a stable dispersion of DPPC liposomes, we combined the two fabrication methods, thin-film rehydration and ultra-sonication. The former is a facile method to produce liposomes with low input energy as well as excellent stability. The latter is an effective way to transfer large multi-lamellar vesicles to small uni-lamellar vesicles. From the TEM analysis, we could confirm that the uni-lamellar structured DPPC liposomes were obtained (Fig. 1C). The particle size, which was analyzed with dynamic light scattering, ranged from 100 nm to 200 nm (Fig. 1D).
3. Results and discussion
To fabricate L-i-M particles, the DPPC liposomes were mixed with the dispersion fluid containing MPC, BIS, Darocure 1173, glycerin, and water. The outer fluid was the paraffin oil containing a hydrophobic surfactant, Abil EM 90. When these two fluids were met at the entrance of the collection capillary tip, a monodisperse L/W/O emulsion was generated. In typical, the drop size was controlled by changing the ratio of the flow rates between the dispersion fluid and outer fluid [21]. To examine how the presence of the liposomes affected formation of emulsion drops, the L/W drops were produced with the dispersion fluid containing different amounts of liposomes (Fig. 2). Then, the radii of the L/W/O emulsion drops produced with varying the concentration of liposomes were normalized with the orifice radius of the end tip of the exit capillary tube. The normalized emulsion sizes were plotted as a function of the scaled flow rate [22,23]. The result is shown in Fig. 3. We observed that the addition of liposomes into the dispersion fluid had no specific influence on the drop generation process, implying that the liposomes were stably suspended in the aqueous monomer solution without any interactions between them or with the water/oil interface.
3.1. Microfluidic strategy for immobilization of liposomes in microgel particles Drop-based microfluidic technology is indeed useful for generation of monodisperse emulsion drops whose typical diameters range from a few micrometers to hundreds of micrometers. In this study, we have built up a microcapillary-based microfluidic platform that enables production of monodisperse liposome-in-water (L/W) drops in paraffin continuous phase. The device fabricated in this study consisted of two separate capillary tubes, as shown in Fig. 1A. The dispersion fluid contained liposomes in the aqueous monomer mixture. In the region near the end tip of the cylinderical tube, the outer fluid focused the dispersion fluid through the cylinderical tube, which is referred as coaxial jetting. This led to formation of a fluid thread which was eventually broken into drops while maintaining the pinch-off prequency due to the Rayleigh-Plateau instability. The suspending L/W drops were solidified to form microgel particles by photo-polymerization. This
3.2. Generation of monodisperse emulsion drops containing liposomes
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Fig. 2. Generation of L/W/O emulsion drops by coaxial jetting. (A) 0 wt% DPPC liposomes in the dispersion fluid, (B) 0.1 wt% DPPC liposomes in the dispersion fluid. After setting the flow rate of the dispersion fluid (DF) to 200 L/h, the flow rate of the outer fluid (OF) was varied.
3.3. Physical immobilization of liposomes in microgel particles Simple photo-polymerization transferred the L/W/O emulsions to the L-i-M microparticles (Fig. 4A). After washing out the paraffin oil, unreacted monomers, and other additives, the L-i-M particles were suspended in water. We observed that the size of the L-iM particles increased by approximately 1.5 times compared with that of the emulsion drops, since the crosslinked PMPC chains were uncoiled in water. To directly show successful immobilization of liposomes in the PMPC network, we visualized the DPPC liposomes by labelling them with a fluorescent phospholipid, Texas RedDHPE. The liposomes also encapsulated the FITC-dextran (70 kDa, 120 A˚ of the Stokes’ diameter). Then, the labeled L-i-M particles were analyzed using a fluorescence microscope. As shown in Fig. 4B–D, green and red fluorescence colors coming from FITC dextran and Texas Red-DHPE, respectively, could be detected from the L-i-M particles. The hydrodynamic diameter of the DPPC liposomes ranges hundreds of nanometers. While on the other, the mesh size of the PMPC microgel was several nanometers. Considering the difference in the length scale, a reasonable postulation is that the DPPC liposomes were physically immobilized in the PMPC hydrogel network.
3.4. Theoretical determination of the mesh size of L-i-M particles To elucidate how the presence of liposomes changes the network property of microgel particles, we determined the mesh size theoretically by using the Peppas and Merrill equation [23–25]. (¯v/V1 )[ln(1 − v2,S ) + v2,S + v22,S ] 1 2 1/3 = − ¯C ¯n M M v2,S v 1 v2,r − 2 v2,S v 2,r
Fig. 3. Normalized drop sizes of L/W/O emulsions with varying the scaled flow rate (QOF /QDF ) (n > 10). The concentration of DPPC liposomes in the dispersion fluid (DF) was changed. 0 wt% (), 0.05 wt% (䊉), 0.1 wt% (), and 0.15 wt% ().
The unperturbed end-to-end chain distance of PMPC polymer, (¯r02 ) by
1/2
(¯r02 )
1/2
, is estimated in terms of the characteristic ratio, Cn , as shown
=l
2,r
¯ C is the number average molecular weight between crosslinks, M ¯n M is the number average molecular weight of the polymer before the cross-linking, v¯ is the specific volume of PMPC (= 0.773 cm3 /g), V1 is the molar volume of the swelling agent, water (= 18 cm3 /mol), is the Flory interaction parameter (= 0.5), v2,r is the volume fraction of the polymer after the crosslinking, but before the swelling, and v2,S is the volume fraction of the polymer after the equilibrium swelling. By determining the volume of the L-i-M particles in a confined state, a swollen state, and a collaped state, v2,S (= 0.090) and v2,r (= 2.266) were gained.
1/2 1/2
Cn
(2)
˚ Mr is the where, l is the carbon–carbon bond length (= 1.54 A), molecular weight of the repeating unit (294 g mol−1 ). Finally, the mesh size, , can be determined, as follows [26]: −1/3
(1)
¯C 2M Mr
= v2,s (¯r02 )
1/2
(3)
To figure out the theoretical relationship between and Cn , we combined Eqs. (2) and (3) and the dependence of on the Cn was observed (Fig. 5A). In principle, the hydrogel mesh size is tunable by changing the degree of crosslinking: Increasing the concentration of BIS resulted in the smaller mesh size. It was interesting in our observation that incorporation of liposomes in the microgel rather increased the mesh size. To quantitatively examine this, we set the Cn at 5 and the corresponding mesh size was plotted as a function of the concentration of BIS (Fig. 5B) [23,27]. The mesh size of the L-i-M particles was around 80 A˚ which is bigger than that of bare PMPC microgel particles (approximately 50–60 A˚ [23]). This means that the liposomes take the space in the hydrogel phase, thus apparently making the mesh size bigger.
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Fig. 4. (A) A schematic for immobilization of liposomes in a microgel particle through photo-polymerization. (B) An bright-field microscopic image of L-i-M particles. (C) A CLSM image of the L-i-M particles containing FTIC-labelled dextran (70 kDa) in the liposome. (D) A CLSM image of the L-i-M particles containing Texas Red-DHPE in the lipid bilayer.
3.5. Molecular network-mediated release behaviors The underlying mechanism for drug release from a matrix has been established on the basis of molecular diffusion [28,29]. The Li-M particles fabricated in this study are attractive in the aspect of controlled drug releasing, since the DPPC liposomes are protected with the PMPC hydrogel matrix. To evaluate the applicability of the
L-i-M system to drug releasing, a hydrophobic model drug, pyrene, was loaded in the DPPC liposomes and then they were physically immobilized in the PMPC microgel particles. The release fashion of pyrene was examined in comparison with the case of directly releasing from the liposomes (Fig. 6). The release experiment was carried out at 42 ◦ C. In this temperature region, the gel-to-liquid crystalline transition of the DPPC bilayer readily occurs, which
Fig. 5. (A) Mesh size versus characteristic ratio for the L-i-M microparticles. The mesh size was tuned with varying the degree of crosslinking. (a) 3 wt% BIS, (b) 6 wt% BIS, and (c) 10 wt% BIS. For comparison, bare microgel particles crosslinked with 3 wt% BIS were also observed (a ). (B) Dependence of the mesh size on the concentration of BIS (n > 10). The open square corresponds to the bare microgel particles.
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Acknowledgement This work was supported by Basic Science Research Program through the National Research Foundation of Korea (NRF) funded by the Ministry of Science, ICT and Future Planning (No. 20080061891) and also by a grant of the Korea Healthcare Technology R&D Project, Ministry of Health & Welfare, Republic of Korea (Grant No.: A103017). This research was also supported by the Korea Institute of Geoscience and Mineral Resources (KIGAM) funded by the Ministry of Knowledge Economy (MEK) of Korea and supported by the R&BD (Research & Business Development) program funded by the Ministry of Trade, Industry and Energy. References Fig. 6. Release profiles of pyrene from the L-i-M microparticles. Liposomes only () and L-i-M microparticles crosslinked with 1.5 wt% BIS (䊉) and 3 wt% BIS (). The concetration of pyrene was set to 0.02 w/v% againt total mass of suspensions (n = 6).
leads to favorable diffusion of pyrene penetrants through the disordered lipid membrane [30,31]. The release of pyrene loaded in the DPPC liposomes showed a typical exponential increase. By contrast, the release amount of pyrene from the L-i-M particles exponentially increased to initial 5 h and then stayed constant. Moreover, the immobilization of the pyrene-loaded liposomes in the microgel particles showed the much lowered cumulative release. This seemed to be attributed to the limited diffusion of pyrene by the enhanced matrix viscosity [32]. These results indicate that the release of pyrene molecules encapsulated in the liposomes was blocked by the surrounding hydrogel matrix, which provides a useful means to improve encapsulation efficiency and to achieve sustained release. 4. Conclusion In summary, we introduce a useful method to improve the structural stability of liposomes by immobilizing them in monodisperse microgel microparticles. The monodisperse L/W/O emulsion with tunable sizes were generated in a capillary-based microfluidic device. Photo-polymerization of the suspending L/W drops allowed us to lock the liposomes in the hydrogel network. The hydrogel mesh property after hybridization with liposomes was investigated by using the Peppas–Merrill equation. We found that the presence of liposomes in the microgel particles increased the apparent mesh size. Further study on drug release revealed that the L-i-M particles could slow down the released rate, which was quite comparable to the case of using liposomes only. The results obtained in this study highlight that hybridization of liposomes with microgel particles enables control over diffusion of drug molecules encapsulated in the liposomes while maintaining their structural stability.
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Contents lists available at ScienceDirect
Colloids and Surfaces B: Biointerfaces journal homepage: www.elsevier.com/locate/colsurfb
Fabrication of monodisperse liposomes-in-microgel hybrid microparticles in capillary-based microfluidic devices Eun Seon Jeong a,1 , Han Am Son b,1 , Min Kyung Kim a , Kyoung-Ho Park c,∗∗ , Sechan Kay d , Pil Seok Chae a , Jin Woong Kim a,c,∗ a
Department of Bionano Technology, Hanyang University, Ansan, Gyeonggi-do 426-791, Republic of Korea Korea Institute of Geoscience and Mineral Resources, Daejeon 305-350, Republic of Korea c Department of Applied Chemistry, Hanyang University, Ansan, Gyeonggi-do 426-791, Republic of Korea d Hwa Costech Co., Seongnam, Gyeonggi-do 463-863, Republic of Korea b
a r t i c l e
i n f o
Article history: Received 4 June 2014 Received in revised form 25 August 2014 Accepted 17 September 2014 Available online 28 September 2014 Keywords: Liposome Microgel Microfluidics Hydrogel mesh size Drug releasing
a b s t r a c t This study introduces a drop-based microfluidic approach to physically immobilize liposomes in microgel (liposomes-in-microgel) particles. For this, we generate a uniform liposomes-in-water-in-oil emulsion in a capillary-based microfluidic device. Basically, we have investigated how the flow rate and flow composition affect generation of emulsion precursor drops in micro-channels. Then, the precursor emulsion drops are solidified by photo-polymerization. From characterization of hydrogel mesh sizes, we have figured out that the mesh size of the liposomes-in-microgel particles is bigger than that of bare microgel particles, since liposomes take space in the hydrogel phase. In our further study on drug releasing, we have observed that immobilization of liposomes in the microgel particles can not only remarkably retard drug releasing, but also enables a sustained release, which stems from the enhanced matrix viscosity of the surrounding hydrogel phase. © 2014 Elsevier B.V. All rights reserved.
1. Introduction The molecular geometry of the unit amphiphiles, which is typically determined by the packing parameter (P = V/aL), can be correlated with their assembled architecture: P is the packing parameter, V and L are the effective volume and length of the amphiphile hydrophobic chain, and a is the amphiphile optimal area at the interface [1,2]. For a packing parameter of ∼1/3, a micelle structure is generated. A bilayered vesicle is made, when the packing parameter is 1/2–1. For a packing parameter of ∼1, a planar bilayer, commonly referred to as lamellar, is formed. The lipid is a sort of amphiphile molecules having both hydrophilic and hydrophobic properties. The packing parameter of lipids ranges 1/2–1, so that they form a lipid membrane with different
∗ Corresponding author at: Department of Applied Chemistry, Hanyang University, 55 Hanyangdaehak-ro, Sangnok-gu, Ansan, Gyeonggi-do 426-791, Republic of Korea. Tel.: +82 31 400 5499. ∗∗ Corresponding author. Tel.: +82 31 400 5409. E-mail addresses: [email protected] (K.-H. Park), [email protected] (J.W. Kim). 1 These authors were equally contributed to this work. http://dx.doi.org/10.1016/j.colsurfb.2014.09.039 0927-7765/© 2014 Elsevier B.V. All rights reserved.
curvatures, which leads to a structurally defined vesicle geometry after the molecular assembly. Liposomes have a bilayered vesicular structure. This structural uniqueness allows them to encapsulate the hydrophilic molecules within the interior aqueous core as well as the hydrophobic molecules within the bilayered lipid shell, respectively [2]. Hence, they are able to protect the encapsulated reactive or sensitive compounds from degradation by shielding and stabilizing against environmental and chemical stresses. Thanks to this function, liposomes are of great interest in pharmaceutical, cosmetic, food and biomedical industries [3,4]. Their size, surface properties, hydrophobicity, and functionality can be tailored by tuning the chemistry and composition of phospholipids and by altering environmental conditions, such as solvent type, ionic strength, pH, and temperature [5–7]. Most of all, the abilities to deliver active and labile molecules into cells or tissues either by vesicle fusion or control of membrane permeability have provided them with more practical applications [8–11]. Despite liposomes have such extensive applicability in academy as well as in industries, they have several critical set-backs, which are currently hampering much wider uses. The carriers made with lipid molecules readily fuse into the membrane bilayers. This makes them fade away quickly upon administrating through organs [12]. Also, liposomes have tendency toward being taken away rapidly
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by immune cells in the blood. In order to resolve these problems, a number of studies have been performed. In typical, they have tailored the periphery of liposomes to physically shield the inner part from external stresses: for instance, the structural stability of liposomes can be enhanced by co-assembly of amphiphilic graft copolymers, so that they can afford to stay in the blood much longer than non-modified ones [13]. One more critical issue that limits the application of liposomes in complex formulations is that the liposomal structure is susceptible to the presence of surfactants with low packing parameters. On adding a very small amount of these surfactants, the liposome-to-micelle transition readily occurs [14,15]. To overcome the above-mentioned drawbacks, in this study, we introduce a facile and trouble-shooting approach in which liposomes were incorporated into microgel microparticles and physically immobilized in the hydrogel network. Hydrogels are able to load and release drug molecules through their web-like molecular network. In principle, the diffusion of drug molecules is mainly ruled by the viscosity of the medium phase, which can be understood by the Stokes-Einstein equation, D = kB T/6a, where D is the diffusion coefficient, kB is the Boltzmann constant, T is the absolute temperature, is the viscosity of the medium, and a is the hydrodynamic radius of the particle. This means that the release kinetics is controllable via regulation of matrix viscosity. To provide a rationale for this, we immobilize the liposomes in uniform microgel particles. This particle system is named as liposomes-in-microgel (L-i-M). For this, we produce a uniform-sized liposomes-in-water-in-oil (L/W/O) emulsion by using a capillary-based microfluidic technique and the L/W drops containing the liposomes are solidified by photopolymerization [16,17]. Basically, the hydrogel phase is made of poly (2-methacryloyloxyethyl phosphorylcholine) (PMPC) having an excellent biocompatibility [18]. Finally, we try to demonstrate that our approach to immobilize the liposomes in the hydrogel microparticles enables a matrix-mediated controlled release. 2. Experimental 2.1. Materials 1,2-Dipalmitoyl-sn-glycero-3-phosphocholine (DPPC) was mercifully offered from Doosan Co. (Korea). 2Methacryloyloxyethyl phosphorylcholine (MPC) was kindly supplied from KCI Co. (Korea). Cetyl PEG/PPG-10/1 dimethicone (Abil EM 90, Evonik, Germany) and hexyltrimethoxysilane (TCI, Japan) were used as received. Chloroform was purchased from Deajung (Korea). Isopropanol was purchased from Samcheon (Korea). Paraffin oil, glycerin, N,N -methylenebisacrylamide (BIS), 2-hydroxy-2-methylpropiophenone (Darocure 1173), fluorescein isothiocyanate (FITC), and fluorescein isothiocyanate dextran (FITC-dextran, 70 kDa) were all purchased from Sigma–Aldrich (USA). Texas Red-1,2-dihexadecanoyl-snglycero-3-phosphoethanolamine (Texas Red-DHPE) was bought from Invitrogen (USA). All other chemicals were reagent grades and used without further purification. For all experiments, deionized double distilled water was used. 2.2. Synthesis of DPPC liposomes DPPC liposomes were prepared using the modified thin-film rehydration method [19,20]. First, DPPC was completely dissolved in chloroform and evaporated with a rotary evaporate at 52 ◦ C for 2 h to remove all the traces of the organic solvent. Then, the thin DPPC film appeared on the wall of the round flask was rehydrated with water and applied mild sonication for 2 h at 65 ◦ C in
order to come off the film into water and to form micron-sized multi-lamellar structured liposomes. The concentration of DPPC in the suspension was set to 1 wt%. To further decrease the particle size and fabricate uni-lamellar liposomes, strong probe-sonication was carried out pulsing every 1 s for 10 min with a power of 130 W at room temperature (VCX130, Sonics & materials Inc.). The size of liposomes was characterized with a dynamic light scattering (ELSZ, Otsuka electronics, Japan). The morphology of liposomes in an aqueous phase was observed with a TEM (Energy-Filtering Transmission Electron Microscope, LIPRA120, Carl Zeiss, Germany). The test samples were negatively stained with 1 wt% of uranyl acetate in the aqueous solution. Then, they were dried in air before TEM observation. 2.3. Fabrication of micro-capillary microfluidic devices To fabricate a capillary-based microfluidic devices, first, a round capillary was tapered by heating and pulling a cylindrical glass capillary (outer diameter = 1.0 mm, World Precision Instruments, USA) with a pipette puller (Model P-97, Sutter Instruments, USA). The end tip of the tapered glass capillaries was cut to the designated diameter using a microforge station (Micro Forge MF 830, Narishige, Japan). To prevent any wetting of the trimmed round capillary by the aqueous inner fluid, hydrophobic coating of the round capillary was conducted with 1 wt% of hexyltrimethoxysilane in toluene. For generation of uniform emulsions, a tapered cylindrical capillary was inserted into a square capillary (Atlantic International Technology, USA). The inner diameter of the square capillary was the same as the outer diameter of the round capillary by 1 mm. Each end of the square capillary was fit into with a needle tip and completely glued with epoxy resin. 2.4. Synthesis of L-i-M microparticles To generate emulsions, each fluid was loaded into a glass syringe (Hamilton Gastight, USA) connected with a polyethylene tube (PE5, Scientific Commodities, USA). Then, the tube was connected with the needle installed at each end of the square capillary. The dispersion fluid (DF) was an aqueous monomer solution containing the liposome suspension (0–15 wt%), MPC (15 wt%), BIS (a crosslinker, 0.5–1.5 wt%), Darocure 1173 (a photo-initiator, 0.5 wt%), and glycerin (30 wt%). The outer fluid (OF) consisted of paraffin oil and 2 wt% Abil EM 90. The flow rate of each fluid was precisely controlled with a syringe pump (Pump 11 Elite, Harvard Apparatus, USA). The emulsion drops were collected through the round capillary and observed with a bright-field microscope (Samwon, NSI-100, Korea). Then, the emulsion drops were solidified by photo-polymerization under UV 365 nm for 1 min. The paraffin oil, remnant monomers, and other additives were thoroughly removed with a large amount of isopropanol by repeated centrifugation at 4000 rpm. Finally, the particles were re-dispersed in water. The volumes of the microgel particles in a confined state, a swollen state, and a collapsed state were determined by measuring their diameters after drying, before swelling, and after swelling with a bright-field microscope. To demonstrate that the liposomes were locked in the microgel particles, we encapsulated 50 L of FITC-dextran in the core and coassembled a small amount of Texas Red-DHPE with the DPPC layers, respectively. The distribution of fluorescence probe molecules in the particles was detected with an inverted fluorescence microscope (IX-81, Olympus, Japan). 2.5. Investigation of releasing profiles Release studies were carried out by using pyrene (ex 336/em 393) as a model drug. For this, a pyrene-loaded liposomal suspension was produced using the same preparation procedure.
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Fig. 1. (A) A schematic of microfluidic set-up for generation of an L/W/O emulsion. (B) A schematic of a liposome particle. (C) A TEM image of the liposome made with 1 wt% DPPC. (D) A particle size distribution of the DPPC liposome.
Undissolved pyrene in the continuous aqueous solution was filtered out with a syringe filter (800 m pores). Pyrene-loaded L-i-M particles were also prepared by following the established method. The concentration of pyrene in both suspensions was regulated with 0.02 w/v%. The suspension was sealed in a dialysis membrane bag (molecular weight cut-off ∼ 10,000, Spectrum lab, Inc., USA) that was immersed into a conical tube containing 40 mL water. Then, the conical tube was gently tumbled with a mixer (RW20D, IKA, Germany) at 42 ◦ C. 150 L of the outer aqueous phase was sampled with the time interval. After each sampling, the outer phase was replaced with the same volume of water. The concentration of pyrene released out was monitored by detecting the intensity of fluorescence with a fluorescence spectroscope (RF-5301PC, Shimadzu, Japan).
process physically fixed the liposomes in hydrogel matrix. The phospholipid used to prepare the liposomes in this study was DPPC (Fig. 1B). To obtain a stable dispersion of DPPC liposomes, we combined the two fabrication methods, thin-film rehydration and ultra-sonication. The former is a facile method to produce liposomes with low input energy as well as excellent stability. The latter is an effective way to transfer large multi-lamellar vesicles to small uni-lamellar vesicles. From the TEM analysis, we could confirm that the uni-lamellar structured DPPC liposomes were obtained (Fig. 1C). The particle size, which was analyzed with dynamic light scattering, ranged from 100 nm to 200 nm (Fig. 1D).
3. Results and discussion
To fabricate L-i-M particles, the DPPC liposomes were mixed with the dispersion fluid containing MPC, BIS, Darocure 1173, glycerin, and water. The outer fluid was the paraffin oil containing a hydrophobic surfactant, Abil EM 90. When these two fluids were met at the entrance of the collection capillary tip, a monodisperse L/W/O emulsion was generated. In typical, the drop size was controlled by changing the ratio of the flow rates between the dispersion fluid and outer fluid [21]. To examine how the presence of the liposomes affected formation of emulsion drops, the L/W drops were produced with the dispersion fluid containing different amounts of liposomes (Fig. 2). Then, the radii of the L/W/O emulsion drops produced with varying the concentration of liposomes were normalized with the orifice radius of the end tip of the exit capillary tube. The normalized emulsion sizes were plotted as a function of the scaled flow rate [22,23]. The result is shown in Fig. 3. We observed that the addition of liposomes into the dispersion fluid had no specific influence on the drop generation process, implying that the liposomes were stably suspended in the aqueous monomer solution without any interactions between them or with the water/oil interface.
3.1. Microfluidic strategy for immobilization of liposomes in microgel particles Drop-based microfluidic technology is indeed useful for generation of monodisperse emulsion drops whose typical diameters range from a few micrometers to hundreds of micrometers. In this study, we have built up a microcapillary-based microfluidic platform that enables production of monodisperse liposome-in-water (L/W) drops in paraffin continuous phase. The device fabricated in this study consisted of two separate capillary tubes, as shown in Fig. 1A. The dispersion fluid contained liposomes in the aqueous monomer mixture. In the region near the end tip of the cylinderical tube, the outer fluid focused the dispersion fluid through the cylinderical tube, which is referred as coaxial jetting. This led to formation of a fluid thread which was eventually broken into drops while maintaining the pinch-off prequency due to the Rayleigh-Plateau instability. The suspending L/W drops were solidified to form microgel particles by photo-polymerization. This
3.2. Generation of monodisperse emulsion drops containing liposomes
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Fig. 2. Generation of L/W/O emulsion drops by coaxial jetting. (A) 0 wt% DPPC liposomes in the dispersion fluid, (B) 0.1 wt% DPPC liposomes in the dispersion fluid. After setting the flow rate of the dispersion fluid (DF) to 200 L/h, the flow rate of the outer fluid (OF) was varied.
3.3. Physical immobilization of liposomes in microgel particles Simple photo-polymerization transferred the L/W/O emulsions to the L-i-M microparticles (Fig. 4A). After washing out the paraffin oil, unreacted monomers, and other additives, the L-i-M particles were suspended in water. We observed that the size of the L-iM particles increased by approximately 1.5 times compared with that of the emulsion drops, since the crosslinked PMPC chains were uncoiled in water. To directly show successful immobilization of liposomes in the PMPC network, we visualized the DPPC liposomes by labelling them with a fluorescent phospholipid, Texas RedDHPE. The liposomes also encapsulated the FITC-dextran (70 kDa, 120 A˚ of the Stokes’ diameter). Then, the labeled L-i-M particles were analyzed using a fluorescence microscope. As shown in Fig. 4B–D, green and red fluorescence colors coming from FITC dextran and Texas Red-DHPE, respectively, could be detected from the L-i-M particles. The hydrodynamic diameter of the DPPC liposomes ranges hundreds of nanometers. While on the other, the mesh size of the PMPC microgel was several nanometers. Considering the difference in the length scale, a reasonable postulation is that the DPPC liposomes were physically immobilized in the PMPC hydrogel network.
3.4. Theoretical determination of the mesh size of L-i-M particles To elucidate how the presence of liposomes changes the network property of microgel particles, we determined the mesh size theoretically by using the Peppas and Merrill equation [23–25]. (¯v/V1 )[ln(1 − v2,S ) + v2,S + v22,S ] 1 2 1/3 = − ¯C ¯n M M v2,S v 1 v2,r − 2 v2,S v 2,r
Fig. 3. Normalized drop sizes of L/W/O emulsions with varying the scaled flow rate (QOF /QDF ) (n > 10). The concentration of DPPC liposomes in the dispersion fluid (DF) was changed. 0 wt% (), 0.05 wt% (䊉), 0.1 wt% (), and 0.15 wt% ().
The unperturbed end-to-end chain distance of PMPC polymer, (¯r02 ) by
1/2
(¯r02 )
1/2
, is estimated in terms of the characteristic ratio, Cn , as shown
=l
2,r
¯ C is the number average molecular weight between crosslinks, M ¯n M is the number average molecular weight of the polymer before the cross-linking, v¯ is the specific volume of PMPC (= 0.773 cm3 /g), V1 is the molar volume of the swelling agent, water (= 18 cm3 /mol), is the Flory interaction parameter (= 0.5), v2,r is the volume fraction of the polymer after the crosslinking, but before the swelling, and v2,S is the volume fraction of the polymer after the equilibrium swelling. By determining the volume of the L-i-M particles in a confined state, a swollen state, and a collaped state, v2,S (= 0.090) and v2,r (= 2.266) were gained.
1/2 1/2
Cn
(2)
˚ Mr is the where, l is the carbon–carbon bond length (= 1.54 A), molecular weight of the repeating unit (294 g mol−1 ). Finally, the mesh size, , can be determined, as follows [26]: −1/3
(1)
¯C 2M Mr
= v2,s (¯r02 )
1/2
(3)
To figure out the theoretical relationship between and Cn , we combined Eqs. (2) and (3) and the dependence of on the Cn was observed (Fig. 5A). In principle, the hydrogel mesh size is tunable by changing the degree of crosslinking: Increasing the concentration of BIS resulted in the smaller mesh size. It was interesting in our observation that incorporation of liposomes in the microgel rather increased the mesh size. To quantitatively examine this, we set the Cn at 5 and the corresponding mesh size was plotted as a function of the concentration of BIS (Fig. 5B) [23,27]. The mesh size of the L-i-M particles was around 80 A˚ which is bigger than that of bare PMPC microgel particles (approximately 50–60 A˚ [23]). This means that the liposomes take the space in the hydrogel phase, thus apparently making the mesh size bigger.
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Fig. 4. (A) A schematic for immobilization of liposomes in a microgel particle through photo-polymerization. (B) An bright-field microscopic image of L-i-M particles. (C) A CLSM image of the L-i-M particles containing FTIC-labelled dextran (70 kDa) in the liposome. (D) A CLSM image of the L-i-M particles containing Texas Red-DHPE in the lipid bilayer.
3.5. Molecular network-mediated release behaviors The underlying mechanism for drug release from a matrix has been established on the basis of molecular diffusion [28,29]. The Li-M particles fabricated in this study are attractive in the aspect of controlled drug releasing, since the DPPC liposomes are protected with the PMPC hydrogel matrix. To evaluate the applicability of the
L-i-M system to drug releasing, a hydrophobic model drug, pyrene, was loaded in the DPPC liposomes and then they were physically immobilized in the PMPC microgel particles. The release fashion of pyrene was examined in comparison with the case of directly releasing from the liposomes (Fig. 6). The release experiment was carried out at 42 ◦ C. In this temperature region, the gel-to-liquid crystalline transition of the DPPC bilayer readily occurs, which
Fig. 5. (A) Mesh size versus characteristic ratio for the L-i-M microparticles. The mesh size was tuned with varying the degree of crosslinking. (a) 3 wt% BIS, (b) 6 wt% BIS, and (c) 10 wt% BIS. For comparison, bare microgel particles crosslinked with 3 wt% BIS were also observed (a ). (B) Dependence of the mesh size on the concentration of BIS (n > 10). The open square corresponds to the bare microgel particles.
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Acknowledgement This work was supported by Basic Science Research Program through the National Research Foundation of Korea (NRF) funded by the Ministry of Science, ICT and Future Planning (No. 20080061891) and also by a grant of the Korea Healthcare Technology R&D Project, Ministry of Health & Welfare, Republic of Korea (Grant No.: A103017). This research was also supported by the Korea Institute of Geoscience and Mineral Resources (KIGAM) funded by the Ministry of Knowledge Economy (MEK) of Korea and supported by the R&BD (Research & Business Development) program funded by the Ministry of Trade, Industry and Energy. References Fig. 6. Release profiles of pyrene from the L-i-M microparticles. Liposomes only () and L-i-M microparticles crosslinked with 1.5 wt% BIS (䊉) and 3 wt% BIS (). The concetration of pyrene was set to 0.02 w/v% againt total mass of suspensions (n = 6).
leads to favorable diffusion of pyrene penetrants through the disordered lipid membrane [30,31]. The release of pyrene loaded in the DPPC liposomes showed a typical exponential increase. By contrast, the release amount of pyrene from the L-i-M particles exponentially increased to initial 5 h and then stayed constant. Moreover, the immobilization of the pyrene-loaded liposomes in the microgel particles showed the much lowered cumulative release. This seemed to be attributed to the limited diffusion of pyrene by the enhanced matrix viscosity [32]. These results indicate that the release of pyrene molecules encapsulated in the liposomes was blocked by the surrounding hydrogel matrix, which provides a useful means to improve encapsulation efficiency and to achieve sustained release. 4. Conclusion In summary, we introduce a useful method to improve the structural stability of liposomes by immobilizing them in monodisperse microgel microparticles. The monodisperse L/W/O emulsion with tunable sizes were generated in a capillary-based microfluidic device. Photo-polymerization of the suspending L/W drops allowed us to lock the liposomes in the hydrogel network. The hydrogel mesh property after hybridization with liposomes was investigated by using the Peppas–Merrill equation. We found that the presence of liposomes in the microgel particles increased the apparent mesh size. Further study on drug release revealed that the L-i-M particles could slow down the released rate, which was quite comparable to the case of using liposomes only. The results obtained in this study highlight that hybridization of liposomes with microgel particles enables control over diffusion of drug molecules encapsulated in the liposomes while maintaining their structural stability.
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