Skip to main content
NCBI home page
NIHPA Author Manuscripts logoLink to NIHPA Author Manuscripts
. Author manuscript; available in PMC: 2011 Apr 12.
Published in final edited form as: Nanotechnology. 2008 May 7;19(18):185102. doi: 10.1088/0957-4484/19/18/185102

New mechanisms for non-porative ultrasound stimulation of cargo delivery to cell cytosol with targeted perfluorocarbon nanoparticles

NR Soman 1, JN Marsh 1, GM Lanza 1, SA Wickline 1
PMCID: PMC3074498  NIHMSID: NIHMS118641  PMID: 21494419

Abstract

The cell membrane constitutes a major barrier for non-endocytotic intracellular delivery of therapeutic molecules from drug delivery vehicles. Existing approaches to breaching the cell membrane include cavitational ultrasound (with microbubbles), electroporation and cell-penetrating peptides. We report the use of diagnostic ultrasound for intracellular delivery of therapeutic bulky cargo with the use of molecularly targeted liquid perfluorocarbon (PFC) nanoparticles. To demonstrate the concept, we used a lipid with a surrogate polar head group, nanogold-DPPE, incorporated into the nanoparticle lipid monolayer. Melanoma cells were incubated with nanogold particles and this was followed by insonication with continuous wave ultrasound (2.25 MHz, 5 min, 0.6 MPa). Cells not exposed to ultrasound showed gold particles partitioned only in the outer bilayer of the cell membrane with no evidence of the intracellular transit of nanogold. However, the cells exposed to ultrasound exhibited numerous nanogold-DPPE components inside the cell that appeared polarized inside intracellular vesicles demonstrating cellular uptake and trafficking. Further, ultrasound-exposed cells manifested no incorporation of calcein or the release of lactate dehydrogenase. These observations are consistent with a mechanism that suggests that ultrasound is capable of stimulating the intracellular delivery of therapeutic molecules via non-porative mechanisms. Therefore, non-cavitational adjunctive ultrasound offers a novel paradigm in intracellular cargo delivery from PFC nanoparticles.

1. Introduction

Targeted therapeutics with synthetic drug delivery vehicles (e.g. lipid or polymer-based nanoparticles) has traditionally been achieved by the endocytosis of the drug delivery vehicle followed by the release of the drug into the cytoplasm. However, lysosomal degradation and inactivation of the drug through this pathway may attenuate therapeutic benefit. Various strategies have been proposed to bypass the lysosomal pathway and achieve non-endocytotic intracellular delivery of therapeutic compounds. These include the conjugation of therapeutic cargo to a cell-penetrating peptide like TAT (Rinne et al 2007, Kerkis et al 2006, McCray et al 2007), electroporation (Fei et al 2007, Ionescu-Zanetti et al 2008, Deora et al 2007, Kusumanto et al 2007), and therapeutic ultrasound with microbubbles (Guzman et al 2001a, 2001b, Taniyama et al 2002, Feinstein 2004, Schlicher et al 2006). A common thread among these strategies is the breakdown of the cell membrane barrier through the formation of pores which allows access to the intracellular compartment.

Ultrasound has been used in conjunction with microbubbles to drive drugs and genes into the cell by causing violent bubble cavitation and jet streaming, resulting in transient pore formation in cell membranes that may be followed by rapid annealing if the membrane damage is not severe. However, as shown both in vitro (Soman et al 2006) and in vivo (Miller and Quddus 2000, Wible et al 2002, Price et al 1998), cavitation also can lead to massive tissue damage and intravascular hemorrhage depending on the energies applied. An unintended consequence may be permanent damage to cells and tissues, calling into question the ultimate clinical utility of microbubbles for drug and gene delivery (Feril et al 2003, ter Haar 2002).

Our laboratory introduced the concept of ‘contact facilitated drug delivery’ with an alternative echogenic nanoparticle contrast formulation that comprised a targeted lipid-coated liquid perfluorocarbon core that was susceptible to activation by ultrasound radiation forces for augmentation of drug delivery to cells by orders of magnitude (Lanza et al 2002, Crowder et al 2005, Hughes et al 2005, Schmieder et al 2005). Drug delivery from these nanoparticles operates through a non-endocytotic energy requiring lipid-raft dependent pathway (unpublished data) whereby the transport of lipids and therapeutic molecules into cells is interlinked. This mode of lipid and drug delivery depends on the biochemical characteristics of the lipid including overall charge, size of the polar head groups, and length of fatty acid side chains (Struck and Pagano 1980, McLean and Phillipis 1981, Gummadi and Kumar 2005). However, whether this mechanism could be used to transport bulky particulates such as nanoparticles into the cytoplasm without disrupting the plasma membrane is not known. This approach could be important for targeted drug delivery to selected cells without the need for adjunctive cell-penetrating peptides such as TAT, which can compromise cell viability (Cardozo et al 2007).

In the present work, we report the use of ultrasound applied at mild diagnostic power levels and frequencies to achieve non-endocytotic delivery of a bulky cargo to the cytoplasm with the use of liquid perfluorocarbon nanoparticles carriers. To demonstrate the concept, we incorporated nanogold conjugated to dipalmitoyl phosphatidyl ethanolamine (nanogold-DPPE), a lipid with bulky polar head group into the lipid monolayer of the PFC nanoparticles as a cargo. In view of an earlier study that demonstrated that nanogold-DPPE would only enter cells in the presence of the pore forming antifungal compound amphotericin (Adler-Moore 1994), we sought to devise a non-cavitational ultrasound approach to achieve transcellular delivery of bulky therapeutic cargo (e.g. nanogold-DPPE) to cells without the need for membrane disruption or the use of adjunctive cell-penetrating peptides.

2. Materials and methods

2.1. Cells and cell-culture

The strategy employed was to target cancer cells with a PFC nanoparticle loaded with nanogold that also contained a specific targeting ligand to avb3 integrins that are expressed on the cell surface. Human C-32 melanoma cells (Washington University Tissue Support Center, St Louis, MO) were grown in minimum essential medium (Earle's) supplemented with 10% bovine fetal bovine serum (Sigma, St Louis, MO). The expression of the integrin receptor avb3 on C-32 melanoma cells was determined by flow cytometry using quantum simply cellular beads (Bangs Laboratories Inc, Fisher, IN) and was in the range of 100 000–150 000 per cell. The cells were grown on fibronectin-coated cell-culture inserts (BD Bioscience, San Jose, CA).

2.2. Synthesis of molecularly targeted perfluorocarbon nanoparticles

Perfluorocarbon nanoparticles targeted to integrin avb3 were synthesized as described earlier (Winter et al 2006). Briefly, a lipid surfactant co-mixture of 99 mol% egg lecithin, 0.1 mol% peptidomimetic vitronectin antagonist specific for avb3 integrins as shown previously (US Patent 6 322 770) conjugated to polyethylene glycol (PEG)2000-phosphatidylethanolamine (Avanti Polar Lipids Inc, Alabaster, AL) and 0.172 mol% nanogold-dipalmitoyl-phosphatidylethanolamine (Nanoprobes Inc, Yaphank, NY) was dissolved in chloroform, evaporated under reduced pressure, dried in a 50 °C vacuum oven and dispersed into water by sonication. The suspension was combined with perfluoro-octylbromide (PFOB) (Gateway Speciality Chemicals, St Peters, MO), and distilled deionized water and continuously processed at 20 000 lbf in−2 for 4 min with an S110 Microfluidics emulsifier (Microfluidics, Newton, MA). This corresponded to 661 nanogold-DPPE molecules per perfluorocarbon nanoparticle. With the use of a laser light-scattering submicron particle size analyzer (Malvern Instruments, Southborough, MA), particle sizes were determined in triplicate at 37 °C to be nominally around 268 nm (range 100–800 nm).

2.3. Ultrasound apparatus

A flat single element piezoelectric transducer with a diameter of 1 cm (natural focus 1.5 cm at 2.25 MHz) was introduced in the cell-culture insert such that the distance between the cell monolayer and the bottom of the transducer was 0.5 cm (figure 1). A 100 MHz synthesized arbitrary waveform generator (Model 395, Wavetek) was used to generate sine wave pulses and the output amplified by using a RF power amplifier (Model 325, LA, Electronic Navigation Industries). Exposure to cells was carried out for approximately 5 min with continuous wave insonification at 2.25 MHz (pressure 0.6 MPa).

Figure 1.

Figure 1

Open in a new tab

In vitro sonication setup showing single element piezoelectric transducer introduced in a cell-culture insert with nanogold-DPPE perfluorocarbon nanoparticles bound to a monolayer of C-32 melanoma cells.

2.4. Transmission electron microscopy

Nanogold-DPPE perfluorocarbon nanoparticles targeted to avb3-were incubated with C-32 melanoma cells in cell-culture inserts at 37 °C for 30 min. The cells were then exposed to continuous wave ultrasound (2.25 MHz, 0.6 MPa) for 5 min. After further incubation for 1 h at 37 °C, the cells were prefixed with 2% glutaraldehyde in 0.1 M sodium cacodylate buffer for 90 min at 4 °C. The gold nanoparticles were silver-enhanced by treating the cells with glycine (50 mM in PBC, pH 7.4), rinsing with Hepes buffer (50 mM), pH 5.8 containing sucrose (200 mM), staining with the HQ silver enhancement kit (Nanoprobes Inc, Yaphank, NY) for 3 min, and fixing with sodium thiosulfate (250 mM) in Hepes buffer (20 mM), pH 7.4 containing 200 mM sucrose. Postfixation was carried out with 0.1% osmium tetroxide in cacodylate buffer (100 mM), pH 7.4 for 30 min at 4 °C. After treatment with uranyl acetate (2% for 48 h at 37 °C), the cell-culture insert membranes were embedded in epoxy resin, thin-sectioned (60 nm) and stained with lead citrate for visualization in a Hitachi H-600 transmission electron microscope.

2.5. Confocal microscopy

C-32 melanoma cells were incubated with nanogold perfluorocarbon nanoparticles as described above. For confocal microscopy however, before exposing the cells to ultrasound, the cells were treated with 20 μg ml−1 rhodamine-wheat germ lectin (Vector Laboratories, Burlingame, CA) for 20 min at 4 °C to stain the plasma membrane. Calcein (Sigma, St Louis, MO), a green-fluorescent marker that is impermeant to viable cells and independent of pH between 6.5 and 12, was introduced in the media at 100 μM. After exposing the cells to ultrasound, the cells were washed extensively and fixed with 4% paraformaldehyde for 10 min at room temperature prior to visualization with a Zeiss 510 confocal microscope.

2.6. Lactate dehydrogenase release assay

Lactate dehydrogenase is an intracellular enzyme that is released from the cells after membrane damage. The LDH release assay was performed according to manufacturer's instructions (BioVision Inc, Mountain View, CA). Briefly, C-32 melanoma cells were exposed to ultrasound (as described above) in the absence or presence of avb3-integrin targeted perfluorocarbon nanoparticles or 0.25% v/v Definity® (Bristol Myers Squibb Medical Imaging, MA). The supernatant was then assayed for the presence of LDH. Cells treated with 1% triton-X were used as a positive control.

3. Results

3.1. Nontreated cell response

Electron micrographic images of silver-enhanced nanogold perfluorocarbon nanoparticles interacting with C-32 melanoma cells are shown in figure 2. At a very early stage of interaction (figures 2(A) and (B)), intact perfluorocarbon nanoparticles were observed binding to the cell membrane. Subsequently, at a later stage (2 h), the nanogold-DPPE from the nanoparticle was observed in the outer cell layer (figures 2(C) and (D)). Similar patterns were apparent at 12 h (data not shown). This confirms that although the nanoparticle fuses with the cell membrane as anticipated from prior work (Crowder et al 2005), nanogold-DPPE is not internalized into the cell under simple incubation conditions thereby reinforcing the findings of Adler-Moore (1994).

Figure 2.

Figure 2

Open in a new tab

A lipid with a bulky polar head group (nanogold-DPPE) is not internalized into the cell when delivered from avb3 targeted perfluorocarbon nanoparticles in the absence of ultrasound. Transmission electron micrographs showing the interaction of silver-enhanced nanogold-DPPE perfluorocarbon nanoparticles with C-32 melanoma cells in the absence of ultrasound. The nanogold-DPPE nanoparticles can be seen bound to the cell membrane ((A) and (B)) after 1 h of incubation at 37 °C while after 3 h nanogold-DPPE is seen to have moved to the cell membrane. No internalization of nanogold-DPPE was seen even after 12 h of incubation. The bar corresponds to 400 nm.

3.2. Ultrasound treated cell response

Figure 3 shows C-32 melanoma cells interacting with perfluorocarbon nanoparticles 2 h after exposure to 5 min of continuous wave ultrasound (2.25 MHz). Nanogold now can be seen to be internalized in the cell beyond the cell membrane with no apparent damage to the cell. The nanogold material appears to traffic in the cytoplasm within uni-vesicular (figures 3(A)–(C)) or multi-vesicular bodies (figure 3(D)).

Figure 3.

Figure 3

Open in a new tab

Ultrasound triggers internalization of nanogold-DPPE delivered from avb3-targeted perfluorocarbon nanoparticles. Transmission electron micrographs of silver-enhanced nanogold-DPPE perfluorocarbon nanoparticles interacting with C-32 melanoma cells after exposure to continuous wave ultrasound (2.25 MHz, 5 min). Nanogold-DPPE can be seen polarized inside intracellular vesicles. The bar corresponds to 400 nm.

3.3. Cell viability

Calcein is an extracellular marker that does not enter intact viable cells. When C-32 melanoma cells were exposed to ultrasound in the presence or absence of nanogold perfluorocarbon nanoparticles, calcein could not be detected in the cell (figure 4(A)). As a positive control, the cells were treated with 0.1% triton-X 100 (Sigma, St Louis, MO), a chemical that is known to permeabilize the cell membrane. For the LDH assay too, cells treated with 0.1% Triton-X 100 were considered as positive control (100% release). The LDH release from cells not treated with ultrasound or PFC nanoparticles was 8.9 ± 4.2%. No significant difference in LDH release was observed after C-32 melanoma cells were exposed to ultrasound in the presence (12.1 ± 4.6%) or absence (11.5 ± 5.4%) of avb3-targeted perfluorocarbon nanoparticles (figure 4(B)). Under similar conditions, cells exposed to ultrasound in the presence of Definity® showed a significant difference (p < 0.05, t-test) in LDH release (24.7 ± 6.4%).

Figure 4.

Figure 4

Open in a new tab

Ultrasound in conjunction with liquid perfluorocarbon nanoparticles does not cause cell membrane damage. (A) Confocal micrographs of C-32 melanoma cells exposed to continuous wave ultrasound (2.25 MHz for 5 min) in the presence or absence of perfluorocarbon nanoparticles. Cells treated with 0.1% triton-X are shown as positive control (red: rhodamine-wheat germ lectin, green: calcein); the bar indicates 20 μm. (B) Release of lactate dehydrogenase (LDH) from C-32 melanoma cells treated with either Definity® or avb3-targeted nanoparticles and ultrasound. Cells treated with triton-X were considered as 100% release. US: ultrasound, NP: nanoparticles.

4. Discussion

Targeted pharmaceutical delivery vehicles typically depend on cellular endocytosis to achieve intracellular delivery of drugs and genes. However, endocytosis of carriers is size dependent and cell-type dependent. For example, the binding of anti-PECAM immunoconjugates to human endothelial cells leads to endocytosis only if their size is less than 300 nm (Wiewrodt et al 2002). Even after endocytosis of the carrier is achieved, the degradation of the active therapeutic molecule in the endosomal–lysosomal pathway may compromise overall efficacy. Existing strategies for bypassing endocytotic pathways to gain direct cytosolic access require breaching the cell membrane barrier (e.g. cell-penetrating peptides, electroporation, ultrasound with microbubbles), which may damage or kill the targeted cell (Price et al 1998, Miller and Quddus 2000, Cardozo et al 2007).

Targeted lipid-coated perfluorocarbon nanoparticles serve as a effective synthetic delivery vehicle for therapeutic molecules that typically bypasses endocytosis. The delivery of drugs and lipids from these nanoparticles to targeted cells likely occurs through contact-mediated vesicle-cell fusion whereby the lipid monolayer of the nanoparticle fuses with the cell membrane allowing the intermixing of nanoparticle lipids with the cell membrane lipids due to lateral diffusion. Further transport into the cell is dependent on lipid flip-flop from the outer to inner membrane and lipid trafficking via caveolae and lipid rafts. The process is energy requiring, as is also the case for endocytosis, and can be inhibited at cold temperatures or by ATP depletion (Partlow K, Lanza G, Wickline S, unpublished data).

An important consideration for such lipid-based membrane-fusional drug delivery approaches is the fact that lipid transport can be affected by a wide variety of factors, chief being the size of the polar head groups and the lengths of the fatty acid side chains (Gummadi and Kumar 2005). Early studies on fusion of lipid vesicles indicated that lipids conjugated to different fluorophores like rhodamine or fluorescein exhibit different diffusion coefficients (McLean and Phillipis 1981). Further, the size of the lipid also affects it's cellular trafficking. For example, as shown by Adler-Moore (1994), lipids conjugated to nanogold do not internalize by themselves but require additional means (e.g., pores) to gain cellular entry. The present studies of transport of bulky polar head group lipids (nanogold-DPPE) demonstrate that nanogold-DPPE was confined to the cell membrane after fusion and lipid mixing (figure 2), but were not otherwise capable of unassisted intracellular transport. As shown earlier (Adler-Moore 1994), transport of the DPPE conjugated to a bulky nanogold (diameter 1.4 nm) was inhibited, likely because it is unable to flip to the inner membrane leaflet.

We have shown previously that clinical ultrasound energy delivered in vitro to cells in the presence of untargeted or targeted PFC nanoparticles can enhance their interaction with cells thereby dramatically increasing lipid delivery many fold (Crowder et al 2005). Further, the uptake of lipidic molecules was far greater from avb3-targeted PFC nanoparticles than non-targeted ones. We have also shown that ultrasound (0.5 and 0.2 MPa) used with microbubbles (Definity®) can cause focal disruptions in the vasculature but when used with PFC nanoparticles do not cause any damage to endothelial cell integrity (Soman et al 2006). Our goal in this study was to employ similar ultrasound parameters and study its effect on the transport of bulky therapeutic cargo (e.g. nanogold-DPPE) from targeted PFC nanoparticles. When C-32 melanoma cells were exposed to continuous wave ultrasound (2.25 MHz for 5 min), nanogold-DPPE was observed inside intracellular vesicles. Ultrasound augmentation of lipid-particulate transport under quite mild conditions was therefore able to overcome the biophysical barriers to lipid trafficking.

Ultrasound in presence of microbubbles has been used to drive drugs and genes into the cell by formation of transient membrane pores as a direct result of cavitation (Feinstein 2004). Recently, low-frequency ultrasound (24 kHz) has been shown to cause transient pores (analogous to wound healing) without the use of cavitational agents (Schlicher et al 2006). However these effects may under certain conditions compromise cell viability (ter Haar 2002) and cause release of LDH as shown in figure 4(B). When C-32 melanoma cells were exposed to continuous ultrasound in the presence or absence of nanogold perfluorocarbon nanoparticles, calcein (an extracellular marker) was not internalized into the cell as demonstrated in the confocal micrographs (figure 4(A)) and no release of intracellular lactate dehydrogenase was observed. These data further indicate that the internalization of nanogold-DPPE is not due to transient discontinuities in the cell membrane, but more likely is related to the specific binding and close apposition of the nanoparticles to the targeted cell membrane.

Previous studies have shown that the cell adhesion molecules such as integrins (e.g., avb3) respond to mechanical stimuli and alter biochemical and gene expression. The integrins thus form the interface for mechano-biological transduction (Ingber 2006). Ultrasound traveling compressional waves induce a temporal-averaged force on the medium (i.e., radiation force) that has been implicated in bone repair (Wang et al 2001), tactile response (Dalecki et al 1995) and auditory response (Tsirulkinov et al 1988). When such a force acts on a PFC nanoparticle, it is clear that it can augment the interaction between the PFC nanoparticle and the cells, and the same applies to these solid gold nanoparticles which otherwise would not penetrate the cell membrane. Ultrasound radiation forces enhance this interaction and lead to the internalization of therapeutic cargo. The mechanisms responsible for this ultrasound transport effect could be multiple, but clearly are not associated with pore formation (sonoporation), membrane disruption, or cavitation. For example, ultrasound could supply the necessary energy to augment lipid flipping, which typically is provided by ATP-requiring flippases, followed by lipid-raft formation and transport (Holthius and Levine 2005, Wenk 2005). A variety of other possibilities exist, but it is interesting to note that once internalized, the lipid–nanogold particles traffic to intracellular vesicles as part of their membrane components (figure 3). Further work will be required to characterize molecular mechanisms responsible and to assess the ultimate disposition of these types of complexes.

In conclusion, our goal was to demonstrate a unique approach to ultrasound-mediated therapeutic cargo delivery to cell cytosol by PFC nanoparticles without deleterious pore formation. Our study shows that non-cavitational ultrasound used in conjunction with liquid perfluorocarbon nanoparticles can induce the uptake and internalization of lipids (and possibly other therapeutic molecules) that would not otherwise be delivered to the cytosol of target cells due to biophysical constraints. Fluorescent microscopy and LDH assay implies that the mechanism of interaction of ultrasound and targeted perfluorocarbon nanoparticles in influencing lipid delivery is not through the formation of pores or transient discontinuities in the cell membrane. The enhanced drug delivery might however be influenced by specific targeting of nanoparticles which when acted upon by ultrasound radiation forces gain sufficient energy to overcome the biophysical barriers to lipid flip-flop or other energy requiring mechanisms. The implications for clinical augmentation of lipophilic drug delivery with conventional ultrasound imaging devices are pertinent to pharmacodynamic profiling.

Acknowledgments

This work was supported in part by the US National Institutes of Health (Grant HL-42590 and HL-59865) to S A Wickline and the American Heart Association Predoctoral fellowship (0515446Z) to N R Soman.

References

  1. Adler-Moore J. AmBisome targeting to fungal infections. Bone Marrow Transplant. 1994;14:S3–7. Review. [PubMed] [Google Scholar]
  2. Cardozo AK, et al. Cell-permeable peptides induce dose-and length-dependent cytotoxic effects. Biochim Biophys Acta. 2007;1768:2222–34. doi: 10.1016/j.bbamem.2007.06.003. [DOI] [PubMed] [Google Scholar]
  3. Crowder KC, Hughes MS, Marsh JN, Barbeiri AM, Fuhrhop RW, Lanza GM, Wickline SA. Sonic activation of molecularly-targeted nanoparticles accelerates transmembrane lipid delivery to cancer cells through contact-mediated mechanisms: implications for enhanced local drug delivery. Ultrasound Med Biol. 2005;31:693–700. doi: 10.1016/j.ultrasmedbio.2005.07.022. [DOI] [PubMed] [Google Scholar]
  4. Dalecki D, Child SZ, Raeman CH, Carstensen EL. Tactile perception of ultrasound. J Acoust Soc Am. 1995;97:3165–70. doi: 10.1121/1.411877. [DOI] [PubMed] [Google Scholar]
  5. Deora AA, Diaz F, Schreiner R, Rodriguez-Boulan E. Efficient electroporation of DNA and protein into confluent and differentiated epithelial cells in culture. Traffic. 2007;8:1304–12. doi: 10.1111/j.1600-0854.2007.00617.x. [DOI] [PMC free article] [PubMed] [Google Scholar]
  6. Fei Z, Wang S, Xie Y, Henslee BE, Koh CG, Lee LJ. Gene transfection of mammalian cells using membrane sandwich electroporation. Anal Chem. 2007;79:5719–22. doi: 10.1021/ac070482y. [DOI] [PubMed] [Google Scholar]
  7. Feinstein SB. The powerful microbubble: from bench to bedside, from intravascular indicator to therapeutic delivery system, and beyond. Am J Physiol Heart Circ Physiol. 2004;287:H450–7. doi: 10.1152/ajpheart.00134.2004. [DOI] [PubMed] [Google Scholar]
  8. Feril LB, Jr, Kondo T, Zhao QL, Ogawa R, Tachibana K, Kudo N, Fujimoto S, Nakamura S. Enhancement of ultrasound-induced apoptosis and cell lysis by echo-contrast agents. Ultrasound Med Biol. 2003;29:331–7. doi: 10.1016/s0301-5629(02)00700-7. [DOI] [PubMed] [Google Scholar]
  9. Gummadi SN, Kumar KS. The mystery of phospholipid flip-flop in biogenic membranes. Cell Miol Biol Lett. 2005;10:101–21. [PubMed] [Google Scholar]
  10. Guzman HR, Nguyen DX, Khan S, Prausnitz MR. Ultrasound-mediated disruption of cell membranes I. Quantification of molecular uptake and cell viability. J Acoust Soc Am. 2001a;110:588–96. doi: 10.1121/1.1376131. [DOI] [PubMed] [Google Scholar]
  11. Guzman HR, Nguyen DX, Khan S, Prausnitz MR. Ultrasound-mediated disruption of cell membranes II. Heterogenous effects on cells. J Acoust Soc Am. 2001b;110:597–606. doi: 10.1121/1.1376130. [DOI] [PubMed] [Google Scholar]
  12. Holthius JC, Levine TP. Lipid traffic: floppy drives and a superhighway. Nat Rev Mol Cell Biol. 2005;6:209–20. doi: 10.1038/nrm1591. [DOI] [PubMed] [Google Scholar]
  13. Hughes MS, Marsh JN, Hall CS, Fuhrhop RW, Lacy EK, Lanza GM, Wickline SA. Acoustic characterization in whole blood and plasma of site-targeted nanoparticle ultrasound contrast agent for molecular imaging. J Acoust Soc Am. 2005;117:964–72. doi: 10.1121/1.1810251. [DOI] [PubMed] [Google Scholar]
  14. Ingber DE. Cellular mechanotransduction: putting all the pieces together again. FASEB J. 2006;20:811–27. doi: 10.1096/fj.05-5424rev. [DOI] [PubMed] [Google Scholar]
  15. Ionescu-Zanetti C, Blatz A, Khine M. Electrophoresis-assisted single-cell electroporation for efficient intracellular delivery. Biomed Microdevices. 2008;10:113–6. doi: 10.1007/s10544-007-9115-x. [DOI] [PubMed] [Google Scholar]
  16. Kerkis A, Hayashi MA, Yamane T, Kerkis I. Properties of cell penetrating peptides (CPPs) IUBMB Life. 2006;58:7–13. doi: 10.1080/15216540500494508. Review. [DOI] [PubMed] [Google Scholar]
  17. Kusumanto YH, Mulder NH, Dam WA, Losen MH, Meijer C, Hospers GA. Improvement of in vivo transfer of plasmid DNA in muscle: comparison of electroporation versus ultrasound. Drug Deliv. 2007;14:273–7. doi: 10.1080/10717540601098807. [DOI] [PubMed] [Google Scholar]
  18. Lanza GM, Yu X, Winter PM, Abendschein DR, Karukstis K, Scott MJ, Chinen LK, Fuhrhop RW, Scherrer DE, Wickline SA. Targeted antiproliferative drug delivery to vascular smooth muscle cells with a magnetic resonance imaging nanoparticle contrast agent: implications for the rational therapy of restenosis. Circulation. 2002;106:2842–7. doi: 10.1161/01.cir.0000044020.27990.32. [DOI] [PubMed] [Google Scholar]
  19. McCray AN, Ugen KE, Heller R. Enhancement of anti-melanoma activity of a plasmid expressing HIV Vpr delivered through in vivo electroporation. Cancer Bio Ther. 2007;6:1269–75. doi: 10.4161/cbt.6.8.4476. [DOI] [PubMed] [Google Scholar]
  20. McLean LR, Phillipis MC. Mechanism of cholesterol and phosphatidylcholine exchange or transfer between unilamellar vesicles. Biochemistry. 1981;20:2893–900. doi: 10.1021/bi00513a028. [DOI] [PubMed] [Google Scholar]
  21. Miller DL, Quddus J. Diagnostic ultrasound activation of contrast agent gas bodies induces capillary rupture in mice. Proc Natl Acad Sci USA. 2000;97:10179–84. doi: 10.1073/pnas.180294397. [DOI] [PMC free article] [PubMed] [Google Scholar]
  22. Price RJ, Skyba DM, Kaul S, Skalak TC. Delivery of colloidal particles and red blood cells to tissue through microvessel ruptures created by targeted microbubble destruction with ultrasound. Circulation. 1998;98:1264–7. doi: 10.1161/01.cir.98.13.1264. [DOI] [PubMed] [Google Scholar]
  23. Rinne J, Albarran B, Jylhävä J, Ihalainen T, Kankaanpää P, Hytönen V, Stayton P, Kulomaa M, Vihinen-Ranta M. Internalization of novel non-viral vector TAT-streptavidin into human cells. BMC Biotechnol. 2007;7:1. doi: 10.1186/1472-6750-7-1. [DOI] [PMC free article] [PubMed] [Google Scholar]
  24. Schlicher RK, Radhakrishna H, Tolentino TP, Apkarian RP, Zarnitsyn V, Prausnitz MR. Mechanism of intracellular delivery by acoustic cavitation. Ultrasound Med Biol. 2006;32:915–24. doi: 10.1016/j.ultrasmedbio.2006.02.1416. [DOI] [PubMed] [Google Scholar]
  25. Schmieder AH, et al. Molecular MR imaging of melanoma angiogenesis with alphavbeta3-targeted paramagnetic nanoparticles. Magn Reson Med. 2005;53:621–7. doi: 10.1002/mrm.20391. [DOI] [PubMed] [Google Scholar]
  26. Soman NR, Marsh JN, Hughes MS, Lanza GM, Wickline SA. Acoustic activation of targeted liquid perfluorocarbon nanoparticles does not compromise endothelial integrity. IEEE Trans Nanobiosci. 2006;5:69–75. doi: 10.1109/tnb.2006.875052. [DOI] [PubMed] [Google Scholar]
  27. Struck DK, Pagano RE. Insertion of fluorescent phospholipids into the plasma membrane of a mammalian cell. J Biol Chem. 1980;255:5404–10. [PubMed] [Google Scholar]
  28. Taniyama Y, Tachibana K, Hiraoka K, Namba T, Yamasaki K, Hashiya N, Aoki M, Ogihara T, Yasufumi K, Morishita R. Local delivery of plasmid DNA into rat carotid artery using ultrasound. Circulation. 2002;105:1233–9. doi: 10.1161/hc1002.105228. [DOI] [PubMed] [Google Scholar]
  29. ter Haar GR. Ultrasonic contrast agents: safety considerations reviewed. Eur J Radiol. 2002;41:217–21. doi: 10.1016/s0720-048x(01)00456-9. [DOI] [PubMed] [Google Scholar]
  30. Tsirulkinov EM, Vartanyan IA, Gersuni GV, Rosenblyum AS, Pudov VI, Gavrilov LR. Use of amplitude-modulated focused ultrasound for diagnosis for hearing disorders. Ultrasound Med Biol. 1988;14:277–85. doi: 10.1016/0301-5629(88)90093-2. [DOI] [PubMed] [Google Scholar]
  31. Wang C, Chen H, Chen C, Yang K. Treatment of nonunions of long bone fractures with shock waves. Clin Orthop. 2001;387:95–101. doi: 10.1097/00003086-200106000-00013. [DOI] [PubMed] [Google Scholar]
  32. Wenk MR. The emerging field of lipidomics. Nature. 2005;4:594–610. doi: 10.1038/nrd1776. [DOI] [PubMed] [Google Scholar]
  33. Wible JH, Jr, Galen KP, Wojdyla JK, Hughes MS, Klibanov AL, Brandenburger GH. Microbubbles induce renal hemorrhage when exposed to diagnostic ultrasound in anesthetized rats. Ultrasound Med Biol. 2002;28:1535–46. doi: 10.1016/s0301-5629(02)00651-8. [DOI] [PubMed] [Google Scholar]
  34. Wiewrodt R, Thomas AP, Cipelletti L, Christofidou-Solomidou M, Weitz DA, Feinstein SI, Schaffer D, Albelda SM, Koval M, Muzykantov VR. Size-dependent intracellular immunotargeting of therapeutic cargoes into endothelial cells. Blood. 2002;99:912–22. doi: 10.1182/blood.v99.3.912. [DOI] [PubMed] [Google Scholar]
  35. Winter PM, et al. Endothelial alpha(v)beta3 integrin-targeted fumagillin nanoparticles inhibit angiogenesis in atherosclerosis. Arterioscler Thromb Vasc Biol. 2006;26:2103–9. doi: 10.1161/01.ATV.0000235724.11299.76. [DOI] [PubMed] [Google Scholar]

RESOURCES