pubs.acs.org/Langmuir © 2010 American Chemical Society
Microfluidic Assembly of Monodisperse, Nanoparticle-Incorporated Perfluorocarbon Microbubbles for Medical Imaging and Therapy Minseok Seo,†,‡ Ivan Gorelikov,‡ Ross Williams,‡ and Naomi Matsuura*,†,‡ †
Department of Medical Biophysics, University of Toronto, and ‡Imaging Research, Sunnybrook Health Sciences Centre, 2075 Bayview Avenue, Toronto, Ontario, Canada M4N 3M5 Received June 4, 2010. Revised Manuscript Received July 13, 2010
New medical imaging contrast agents that permit multiple imaging and therapy applications using a single agent can result in more accurate diagnosis and local treatment of diseased tissue. Solid nanoparticles (NPs) (5-150 nm in size) have emerged as promising imaging and therapy agents, as have micrometer-scale, perfluorocarbon gas-filled microbubbles (MBs) used in patients as intravascular ultrasound contrast agents. We propose that the modular combination of small, solid NPs and larger, highly compressible MBs into a single agent is an effective way to attain the desired complementary and hybrid properties of two very different agents. Presented here is a new strategy for the simple and robust incorporation of various medical NPs with monodisperse MBs based upon the controlled pH-based regulation of the electrostatic attraction between NPs and the MB shell. Using this simple approach, microfluidicgenerated, protein-lipid-coated, perfluorobutane MBs (with size control down to 3 μm) were incorporated with silicacoated NPs, including CdSe/ZnS quantum dots, gold nanorods, iron oxide NPs, and Gd-loaded mesoporous silica NPs. The silica interface permits NP inclusion within MBs to be independent of NP composition, morphology, and size. Significantly, the NP-incorporated MBs (NP-MBs) diluted in saline were detectable using low-pressure ultrasound, and the monodisperse MB platform can be produced at high-throughput, sufficient for in vivo usage (106 MB/sec). The modular synthesis of a variety of NP-MBs can facilitate flexible, user-defined, multifunctional imaging and therapy agents tailored for specific applications and disease types.
Introduction Medical agents that permit multiple imaging or therapy applications using a single agent can provide concurrent contrast imaging for multiple modalities with complementary spatial, temporal, and depth resolution for more accurate diagnosis and local treatment of diseased tissues. For contrast imaging, each imaging modality (e.g., ultrasound, magnetic resonance, photoacoustic, and optical imaging) requires different contrast agents with particular physical and material properties tailored to its particular contrast mechanism. For example, clinical ultrasound imaging contrast agents are typically intravascular, highly echogenic, perfluorocarbon gasfilled, 1-7 μm-sized microbubbles (MBs).1 For other imaging modalities, nanoparticles (NPs), ca. 5-150 nm in size, are often used. For example, quantum dot (QD) NPs can be used for optical imaging, gold nanorods (Au NRs) can be used for photoacoustic imaging, and iron oxide or Gd-loaded NPs can be used for magnetic resonance imaging. Furthermore, some of these same imaging contrast agents can also be used for localized treatment in vivo. In particular, MBs can be disrupted by ultrasound for enhancement of hyperthermia or in vivo delivery of colocalized drugs or genes.2-4 Since the in vivo biodistribution of contrast and therapy agents are size and material dependent, separate injections of different agents will not result in complementary imaging or treatment at the same site. For this reason, the most common approach for *To whom correspondence should be addressed. E-mail: matsuura@ sri.utoronto.ca. (1) Schutt, E. G.; Klein, D. H.; Mattrey, R. M.; Riess, J. G. Angew. Chem., Int. Ed. 2003, 42, 3218–3235. (2) Unger, E. C.; Porter, T.; Culp, W.; Labell, R.; Matsunaga, T.; Zutshi, R. Adv. Drug Delivery Rev. 2004, 56, 1291–1314. (3) Lindner, J. R. Nat. Rev. Drug Discovery 2004, 3, 527–532. (4) Ferrara, K. W. Adv. Drug Delivery Rev. 2008, 60, 1097–1102.
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multifunctional agent development is the modular combination of complementary, individual agents together into a single agent. For applications in conjunction with ultrasound-based imaging and therapy, the combined agent typically consists of a MB platform upon which the smaller imaging and therapy NPs are loaded. Despite the synthetic challenge of adding NPs to MBs, several research groups have successfully demonstrated NP-MB multimodal imaging and therapy (i.e., drug and gene delivery) agents.2,5-11 Ferrara et al.6 demonstrated a model drug delivery system which consisted of avidinated fluorescent nanobeads bound directly to the biotinylated lipid shells of MBs. Gu et al.7-9 synthesized a polymeric MB shell containing Fe3O4 NPs for ultrasound and magnetic resonance imaging. Recently, Dai et al.11 developed a method of loading QDs onto MBs via layer-by-layer deposition of poly(allylamine hydrochloride) for optical and ultrasound imaging. However, in each case, the attachment of NPs to MBs depends on specific binding interactions between different imaging NPs and MBs. Thus, each of these methods is designed for a specific dual-imaging system and cannot be easily translated to other NP-MB systems. Finally, all (5) Price, R. J.; Skyba, D. M.; Kaul, S.; Skalak, T. C. Circulation 1998, 98, 1264– 1267. (6) Lum, A. F. H.; Borden, M. A.; Dayton, P. A.; Kruse, D. E.; Simon, S. I.; Ferrara, K. W. J. Controlled Release 2006, 111, 128–134. (7) Yang, F.; Li, Y. X.; Chen, Z. P.; Zhang, Y.; Wu, J. R.; Gu, N. Biomaterials 2009, 30, 3882–3890. (8) Yang, F.; Gu, A. Y.; Chen, Z. P.; Gu, N.; Ji, M. Mater. Lett. 2008, 62, 121– 124. (9) Yang, F.; Li, L.; Li, Y. X.; Chen, Z. P.; Wu, J. R.; Gu, N. Phys. Med. Biol. 2008, 53, 6129–6141. (10) Stride, E.; Pancholi, K.; Edirisinghe, M. J.; Samarasinghe, S. J. R. Soc., Interface 2008, 5, 807–811. (11) Ke, H.; Xing, Z. W.; Zhao, B.; Wang, J. R.; Liu, J. B.; Guo, C. X.; Yue, X. L.; Liu, S. Q.; Tang, Z. Y.; Dai, Z. F. Nanotechnology 2009, 20, 1-8.
Published on Web 07/28/2010
DOI: 10.1021/la102272d
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previous NP-MB systems are based on polydisperse MB scaffolds, which are less ideal than monodisperse MBs for specific ultrasound imaging applications. MB monodispersity can enhance the sensitivity of ultrasound imaging because the echogenic response and disruption threshold of MBs can be controlled by their size.12,13 The production of MBs with narrow size distributions can result in lower injected agent doses into patients, as only small proportions of traditional polydisperse MB populations are resonant in the ultrasound field. However, narrow size distributions of MBs produced via conventional sonication or mechanical agitation are difficult to achieve due to the nonuniform stresses generated during their synthesis. Recently, progress has been made in the field of microfluidic-based generation of extremely monodisperse MBs.14-16 Generally, monodispersity is defined as having coefficient of variation (CV, the standard deviation divided by the mean diameter 100) of less than 5.0%.17,18 There are significant challenges that must be addressed in the development of NP-incorporated, monodisperse MBs suitable for in vivo ultrasound imaging. First, the MB platform must be synthesized at high-enough yields (107-109 MB/mL concentration) for injection into preclinical models or patients. This requires that MBs are produced at a generation rate of ∼106 MB/sec, which is challenging because of the characteristic, one-by-one production of single MBs by microfluidics. Second, the ideal MB size is determined by its ability to resonate in the applied ultrasound field and its ability to travel through the vasculature. For this reason, MB size is limited to the 1-7 μm range,1 which is close to the stability limit of MB production by microfluidics. Third, the properties of the as-synthesized monodisperse MB must be designed for strong ultrasound response (i.e., thin, highly flexible MB shell), in vivo compatibility (i.e., biodegradable MB shell components), and NP inclusion (i.e., tunable shell surface charge) while retaining enough stability for NP inclusion without destroying fragile MBs or loss of monodispersity. Fourth, NP-MB agents must be stable in vivo for transcardiac and transpulmonary passage and survive in the bloodstream long enough to provide sufficient data collection (i.e., new NP-MB agents must be stable at low concentrations at physiological pH and salinity, without excess surfactant in the dispersed phase).1 This is not trivial, as monodisperse MBs are typically most stable within their native synthesis solution. To further increase MB lifetimes, biocompatible perfluorobutane (PFB) gas, which is less soluble in water and blood than air-filled MBs,1,19 is typically used as the MB core gas material. In this paper, we introduce a simple method to assemble multifunctional, MB-based agents based on the loading of charged medical NPs onto monodisperse, protein-lipid-stabilized MB platforms. Specifically, we outline a new strategy for the simple and robust incorporation of various silica-coated NPs onto (12) Talu, E.; Hettiarachchi, K.; Zhao, S.; Powell, R. L.; Lee, A. P.; Longo, M. L.; Dayton, P. A. Molecular Imaging 2007, 6, 384–392. (13) Hettiarachchi, K.; Lee, A. P.; Zhang, S.; Feingold, S.; Dayton, P. A. Biotechnol. Prog. 2009, 25, 938–945. (14) Garstecki, P.; Gitlin, I.; DiLuzio, W.; Whitesides, G. M.; Kumacheva, E.; Stone, H. A. Appl. Phys. Lett. 2004, 85, 2649–2651. (15) Hettiarachchi, K.; Talu, E.; Longo, M. L.; Dayton, P. A.; Lee, A. P. Lab Chip 2007, 7, 463–468. (16) Lee, M. H.; Prasad, V.; Lee, D. Langmuir 2010, 26, 2227–2230. (17) Prasad, N.; Perumal, J.; Choi, C. H.; Lee, C. S.; Kim, D. P. Adv. Funct. Mater. 2009, 19, 1656–1662. (18) Jillavenkatesa, A.; Dapkunas, S. J.; Lum, L. H. NIST Special Publication 960-1; U.S. Government Printing Office: Washington, DC, 2001; (a particle distribution may be considered monodisperse if at least 90% of the distribution lies within 5% of the median size ). (19) Kabalnov, A.; Klein, D.; Pelura, T.; Schutt, E.; Weers, J. Ultrasound Med. Biol. 1998, 24, 739–749.
13856 DOI: 10.1021/la102272d
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MBs based on controlled pH-based20 regulation of the electrostatic attraction between NPs and monodisperse MBs (with size control down to 3 μm) produced via microfluidics at yields of ∼106 MB/sec. For this work, the MB shell was uniquely formulated to achieve stability at low concentration in saline, for attachment of charged NPs, and for strong resonance in the ultrasound field. The use of silica as a common interface to the charged MB shell allowed the demonstration of modular attachment of various NPs via surface charge. Using this simple approach, we incorporated silica-coated CdSe/ZnS QDs, Au NRs, Fe3O4, and Gd-loaded mesoporous silica NPs, into protein-lipid-coated PFB MBs for potential applications in optical, photoacoustic, and MR imaging, respectively. Incorporation of NPs into MBs was found to be independent of NP size and surface morphology, and NP-MBs were successfully detected using lowpressure ultrasound and an MB specific imaging mode.
Experimental Section Materials. Alginic acid sodium salt (alginate), lysozyme from chicken egg white (lysozyme), and other chemicals were purchased from Sigma-Aldrich. Lipids, 1,2-dipalmitoyl-sn-glycero-3-phosphocholine (DPPC), and N-(carbonyl-methoxy-polyethyleneglycol5000)-1,2-dipalmitoyl-sn-glycero-3-phosphoethanolamine, sodium salt (MPEG-5000-DPPE), were purchased from Bio-Lab Ltd. (Israel). Deionized water (Millipore Milli-Q grade) with resistivity of 18.2 MΩ was used in all the experiments. Poly(dimethylsiloxane) (PDMS, Sylgard 184), and photoresist resin (SU-8 series) were purchased from Dow Corning (USA), and Microchem Co. (MA), respectively. PFB was purchased from FluoroMed, L.P. (TX). All chemicals were used as-received. Methods. Characterization. Electron microscopy images were acquired on a Hitachi S-5200 scanning electron microscope (SEM) and a Hitachi HD-2000 scanning transmission electron microscope (STEM). Absorbance spectra for Au NR and Fe3O4 characterization were obtained using a Cary-6000i UV-vis-NIR spectrophotometer. Zeta (ζ)-potential and hydrodynamic size measurements of the NPs were taken using a Malvern Zeta-Sizer 3000HS instrument. MB concentrations were measured using a Coulter Counter (Multisizer 3, Beckman Coulter, Inc., USA). Elemental composition of silica-coated NPs was determined using an Optima 3000 ICP-AES system. The process of the generation of MBs was imaged using a high speed CCD camera (CoolSNAP HQ2, Photometrics, USA), with exposure times of 1 μs. Fluorescence microscopy of QD-incorporated MBs were analyzed using a Zeiss Axiovert 200 M microscope with Axiovision imaging software and a standard TRITC filter set (i.e., an excitation filter range of 545-565 nm and an emission filter range of 580620 nm). MB size was analyzed using Image Pro Plus (Media Cybernetics, USA) software (>400 MBs). MB size measurements were taken in three locations: immediately after the orifice; before the outlet (2 and 7 cm from the orifice); and after exiting the microfluidic device (after collection in a glass container). A Philips iU22 ultrasound system equipped with an L9-3 probe was used for ultrasound imaging. Preparation of Protein-Lipid Solution. The protein solution was prepared by adding 0.10-0.20 wt % lysozyme and 0.100.20 wt % alginate into a solution of 10 vol % glycerol, 10 vol % propylene glycol, and 80 wt % DI water (GPW) at pH 11.5. The lipid mixture, 0.9 mol % of DPPC, and 0.1 mol % MPEG-5000DPPE were dissolved in chloroform, evaporated under a gentle flow of nitrogen gas, dried in a 50 °C vacuum oven overnight, and then dispersed into the GPW. Freeze-thaw cycles were carried out after preparation of the lipid mixture. The concentration of the lipid mixture was 7.5 mg/mL. Before the experiment, a 9:1 w/w
(20) Beebe, D. J.; Moore, J. S.; Bauer, J. M.; Yu, Q.; Liu, R. H.; Devadoss, C.; Jo, B. H. Nature 2000, 404, 588–590.
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Figure 1. (a) Schematic of the microfluidic device used to generate PFB MBs. (b) Typical optical microscopy image of PFB MBs obtained in the PDMS microfluidic device. Scale bar = 20 μm. mixture of the protein-lipid solution was prepared and sonicated at room temperature for 10 min. Fabrication of Microfluidic Devices. A mask pattern was designed, transferred to a graphic software file (AutoCAD, Autodesk Inc., USA), and printed on a transparency. Laser photoplotted transparencies (20 000 dpi) (CAD/Art Services, Poway, CA) were used as masks to prepare a master in a thin film of photoresist. Photolithographic masters with features of SU-8 25 photoresist (MicroChem Corp., USA) were prepared in bas-relief on silicon wafers. A negative replica of this master in a poly(dimethylsiloxane) (PDMS, Sylgard 184, Dow Corning USA) elastomer was used as the mold for soft lithography.21,22 The height of the microfluidic channels was ∼25 μm.
Results and Discussion 1. Microfluidic Generation of PFB MBs. For the production of monodisperse, perfluorocarbon MBs at high-throughput and small sizes suitable for NP loading and in vivo ultrasound applications, new microfluidic masks were designed. Specifically, a three-inlet microfluidic flow-focusing device was designed to generate MBs of controllable sizes with narrow size distribution (see Supporting Information, Figure S1 for device dimensions). PFB gas, with boiling point, diffusion coefficient, water solubility, and Ostwald coefficient of -2 °C, 6.9 (D 109, m2/s, at 20 °C), 0.021 (mol/m3, at 25 °C), and 202 (L 106, at 25 °C), respectively,19 was used as the gas phase. The dispersed phase (PFB) was introduced in the central channel, while the continuous phase (aqueous shell solution) was supplied to the two side channels of the device (Figure 1). A pressure gradient along the long axis of the device forced the two immiscible fluids (PFB gas and continuous phase) through the narrow orifice of the microfluidic device. The continuous phase surrounded the inner, immiscible PFB gas so that the inner thread became unstable and broke in the orifice in a periodic manner to release MBs into the downstream channel.14 The geometry of the device, gas pressure (PPFB), and flow rate of the continuous phase (Qc) determined the size of PFB MBs. Generation of MBs in the flow-focusing regime produced MBs with diameters comparable to the orifice width.15,23 (21) Xia, Y. N.; Whitesides, G. M. Angew. Chem., Int. Ed. 1998, 37, 551–575. (22) Eddington, D. T.; Puccinelli, J. P.; Beebe, D. J. Sens. Actuators B 2006, 114, 170–172. (23) Seo, M.; Nie, Z. H.; Xu, S. Q.; Mok, M.; Lewis, P. C.; Graham, R.; Kumacheva, E. Langmuir 2005, 21, 11614–11622. (24) Cavalieri, F.; Ashokkumar, M.; Grieser, F.; Caruso, F. Langmuir 2008, 24, 10078–10083.
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For the continuous phase, a protein-lipid solution dispersed in glycerol, propylene glycol, and DI water (GPW) at pH 11.5 was used.24,25 Lysozymes are well-characterized folded proteins and are easily obtained in a purified form.26 Since the adsorption of proteins onto organic/inorganic NPs is well known, strong electrostatic interactions were anticipated between negatively charged NPs and positively charged lysozymes below its isoelectric point (IEP).26-28 Since stabilization of microfluidic-produced PFB MBs by lysozymes and proteins alone requires excess surfactant in the dispersed phase to retain their concentration and monodispersity, we developed a new shell formulation based on a combination of lipids, DPPC, and MPEG-5000-DPPE, which costabilized the PFB MBs with the proteins such that the MBs were stable after dilution in saline. The solution was prepared with GPW to control the viscosity of the continuous solution so that smaller and more stable MBs could be obtained at high-throughputs compared to using a less viscous continuous phase.12,29 The continuous phase was supplied to the device through polyethylene or Teflon tubing attached to a syringe operated by a digitally controlled syringe pump. To achieve highly monodisperse PFB MBs in the aqueous phase, the device material required a high affinity for the continuous phase.22,23 Since PFB is a hydrophobic and lipophobic gas, PFB MBs were successfully generated in the freshly prepared PDMS microfluidic device. To examine the effect of flow rate, Qc was varied from 1.0 to 7.0 mL/h and the resultant MB diameter was measured (Figure 2a-f). The gas pressure (PPFB) was varied from 5 to 22 psi. Because of the relatively narrow width of the orifice, the velocity of the continuous phase in the orifice was high. The mean diameter (Dm) and the CV of PFB MBs stabilized by the protein-lipid solution were plotted as a function of the Qc at a PPFB of 17.5 psi (Figure 2g). For controlled generation of MBs with a narrow size distribution, two necessary conditions of the quasistatic collapse of the thread of the PFB phase were met: the geometry of the thread was stable against the capillary instability and the rate of thinning of the thread was much slower than typical rates of relaxation of the interfacial energy.30-32 Foaming in the microfluidic devices occurred through a series of equilibrium states, each with a minimum interfacial energy at the gas-fluid interface. The volume of the MBs decreased with increasing Qc for a fixed PPFB. Under these conditions, MBs with sizes ranging from 3 to 9 μm and with CV values ranging from 1 to 3% were successfully generated in the flow-focusing regimes,14 although on rare occasions, some MB generation was observed in the thread-forming regime.31 Depending on the combination of PPFB and Qc, MBs smaller in diameter than the dimension of the orifice were successfully generated. Although small MBs (
