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Pancreatic ductal adenocarcinoma (PDA) is the fourth most common cause of cancer death in the United States, with a median survival time of only 3–6 months for forty percent of patients. Current treatments are ineffective and new PDA therapies are urgently needed. In this context, ultrasound-mediated chemotherapy by polymeric micelles and/or nanoemulsion/microbubble encapsulated drugs may offer an innovative approach to PDA treatment. PDA xenografts were orthotopically grown in the pancreas tails of nu/nu mice by surgical insertion of Red Fluorescence Protein (RFP)-transfected MiaPaCa-2 cells. Tumor growth was controlled by fluorescence imaging. Occasional sonographic measurements correlated well with the formal tumor tracking by red fluorescence. Tumor accumulation of paclitaxel-loaded nanoemulsion droplets and droplet-to-bubble transition under therapeutic ultrasound was monitored by diagnostic ultrasound imaging. MiaPaCa-2 tumors manifested resistance to treatment by Gemcitabine (GEM). This drug is the gold standard for PDA therapy. The GEM-resistant tumors proved sensitive to paclitaxel. Among six experimental groups studied, the strongest therapeutic effect was exerted by the following drug formulation: GEM + nanodroplet-encapsulated paclitaxel (nbGEN) combined with tumor-directed 1-MHz ultrasound that was applied for 30 s four to five hours after the systemic drug injection. Ultrasound-mediated PDA therapy by either micellar or nanoemulsion encapsulated paclitaxel resulted in substantial suppression of metastases and ascites suggesting ultrasound-enhanced killing of invasive cancerous cells. However, tumors relapsed after the completion of therapy, indicating survival of some tumor cells. The recurrent tumors manifested development of paclitaxel resistance. Ultrasound imaging suggested non-uniform distribution of nanodroplets in the tumor volume due to irregular vascularization, which may result in the development of zones with sub-therapeutic drug concentration. This is implicated as a possible cause of the resistance development, which may be pertinent to various modes of tumor nanotherapy.
The application of biomedical nanotechnology holds promise for solving major problems in current tumor chemotherapy related to the severe side effects of chemotherapeutic drugs and development of drug resistance. During the last decade, the advances of nanomedicine have allowed for the combination of various functionalities in one molecular or supramolecular construct, i.e. polymeric bioconjugate or nanoparticle. The nanoparticles include polymeric micelles, liposomes, hollow particles, nano- or microemulsion droplets etc. Various chemotherapeutic drugs, imaging agents, and targeting moieties may be encapsulated in the same nanocontainer. This ability to combine chemotherapeutic agents and imaging markers is especially important in oncological practice potentially allowing early assessment of response to treatment.
Among various imaging modalities, ultrasound is the most cost effective. Ultrasound provides real-time information. The lack of ionizing radiation (in contrast to e.g. CT) is also an advantage. Ultrasound imaging may be combined with ultrasound-mediated drug delivery by ultrasound-responsive nanoparticles. This allows double tumor targeting because nanoparticles accumulate in tumor tissue via the enhanced permeability and retention (EPR) effect1, followed by local release of the encapsulated drug in tumor tissue in response to tumor-directed ultrasound2–6. These drug nanocarriers are expected to passively accumulate in tumors because of the enhanced permeability and retention (EPR) effect, which is based on defective tumor microvasculature that allows extravasation of drug-loaded nanoparticles through large inter-endothelial gaps7, 8. In addition, poor lymphatic drainage of tumors results in longer retention of extravasated particles in tumor tissue. In contrast to tumors, blood vessels in normal tissues have tight inter-endothelial junctions which do not allow extravasation of nanoparticles.
Effective tumor accumulation of nanoparticles via the EPR effect requires sufficient particle residence time in circulation; to provide for this, nanoparticles are commonly coated with poly(ethylene oxide) chains that suppress blood protein adsorption and prevent particle recognition by the cells of the reticulo-endothelial system.
Ultrasound as a drug delivery modality offers a number of important advantages in comparison with other physical modalities. Sonication may be performed non-invasively or minimally invasively through intraluminal, laparoscopic or percutaneous means. For extracorporeal sonication, the transducer is placed in contact with a water-based gel or a water layer on the skin, and no insertion or surgery is required. Ultrasound can be directed toward deeply located body sites and tumor sonication with millimeter precision is feasible.
During the last decade, a number of groups have concentrated their efforts on developing ultrasound-responsive drug delivery systems for oncological applications9–18. Ultrasound as a component of a drug delivery modality may be applied with a variety of drug carriers. Development of ultrasound-responsive stable liposomes that manifested prolonged circulation time and effective tumor targeting was recently reported. Ultrasound-induced heating triggered phase transition in the phospholipid membrane and released the drug in the target region. Ultrasound-triggered delivery of paclitaxel (PTX) to the monolayers of prostate cancer cells from a phospholipid-coated perfluorohexane nanoemulsion developed by ImaRx was also reported19.
Polymeric micelles that combined passive tumor-targeting ability with ultrasound responsiveness were suggested in earlier works of Rapoport’s group2, 6, 20–25. It was demonstrated that drug-loaded polymeric micelles accumulated passively in tumor tissue26. Ultrasonic irradiation of the tumor enhanced micelle uptake, triggered drug release from the micelles and transiently altered cell membrane permeability, resulting in effective intracellular drug internalization by tumor cells. Ultrasound irradiation also enhanced drug diffusion throughout tumor volume, thus reducing drug concentration gradients2, 6, 26. Substantial reduction of the tumor growth rate was achieved using this drug delivery modality6, 27.
All therapeutic modalities associated with application of physical stimuli, such as tumor ablation induced by radiofrequency, microwave, or ultrasound as well as local mild hyperthermia require imaging of targeted sites prior to and during treatment. Though various imaging modalities may be used, combining ultrasound-mediated drug delivery with ultrasound imaging appears especially attractive by virtue of offering real time information, versatility, and cost effectiveness. In this context, during the last decade, microbubbles have attracted attention as drug carriers and enhancers of drug and gene delivery. In current clinical practice, microbubbles have been used as ultrasound contrast agents for cardiovascular imaging28. Several research groups have concentrated their efforts on developing microbubble-based drug delivery systems2, 3, 11–13, 15, 16, 19, 29–42.
The most cost-effective way to solve this problem would be to impart drug carrier properties to FDA approved ultrasound contrast agents such as Optison (Amersham Inc.) or Definity (Lanteus Medical Imaging Inc.). However, currently used contrast agents present a number of inherent problems as drug carriers. Their very short circulation time (minutes) and relatively large size (two to ten microns) do not allow effective extravasation into tumor tissue, which is an essential prerequisite for effective drug targeting.
A possible way to solve the above problem may consist in developing drug-loaded, nano-scaled microbubble precursors that would effectively accumulate in tumor tissue and then convert into microbubbles in situ after tumor accumulation. With this in mind, we have recently developed block copolymer stabilized echogenic (i.e. ultrasound contrast generating) perfluoropentane (PFP) nanoemulsions that convert into microbubbles under ultrasound irradiation2, 3, 43, 44. The nanoemulsions are produced from drug-loaded poly(ethylene oxide)-co-poly(L-lactide) (PEG-PLLA) or poly(ethylene oxide-co-polycaprolactone (PEG-PCL) micelles by introducing a phase-shift perfluorocarbon compound, perfluoropentane (PFP). The PFP has a boiling temperature of 29 °C. However, PFP nanoemulsions manifest remarkable thermal stability due to the excessive pressure called Laplace pressure inside nanodroplets4, 43. The thermal stability of PFP nanoemulsions prevents their conversion into microbubbles in circulation. They maintain their nanoscale size, which allows tumor accumulation by extravasation through leaky tumor microvasculature (i.e. passive targeting)2, 4. Without ultrasound, the drug is tightly retained inside the droplet walls, which is important for effective tumor targeting4. After tumor accumulation, droplet-to-bubble conversion is triggered by tumor-directed ultrasound. This effect is called acoustic droplet vaporization (ADV)45. Tumor irradiation by therapeutic ultrasound induces localized drug release from nanodroplets and microbubbles and effective intracellular drug uptake by tumor cells, which in turn, results in effective tumor regression. In addition, primary small microbubbles coalesce into larger microbubbles and provide a long-lasting ultrasonic contrast2–4. Using this treatment modality, effective tumor regression was observed in mouse models of breast and ovarian cancer2–4; the first promising results for pancreatic cancer were presented in ref.4.
Pancreatic ductal adenocarcinoma (PDA) is the fourth most common cause of cancer death in the United States. The annual incidence rate of pancreatic cancer is almost identical to the mortality rate; approximately 37,000 new cases are diagnosed each year in the United States, and approximately 33,000 patients die from this disease. Only 4% of patients are alive 5 years after the time of diagnosis. Most PDA presentations are inoperable at the time of diagnosis due to the extensive tumor burden, local invasion, poor general health, and multiple aggressive micrometastases that are resistant to chemotherapy and radiation treatment. About 40% of patients have a dismal prognosis; median survival time is only 3–6 months46.
The only FDA-approved chemotherapeutic agent for PDA is a nucleoside analogue gemcitabine (GEM), but the partial response rate to chemotherapy is well below 10%47, 48 most probably due to the development of GEM resistance in the course of chemotherapy. Despite low effectiveness, GEM remains the cornerstone of neoadjuvant and adjuvant chemotherapy in pancreatic cancer and imparts a progression-free survival interval ranging from 0.9 to 4.2 months49. Neither biotherapies based on the targeted gene therapy, use of antibodies against Vascular Endothelial Growth factor (VEGF) or Epithelial Growth Factor (EGF-R) receptors50 nor high-intensity ultrasonic tumor ablation51–54 proved successful. An attempt was made to treat pancreatic cancer by extracorporeal high intensity focused ultrasound (HIFU), with only marginal success55.
New approaches to PDA treatment are urgently needed. In this context, ultrasound-mediated chemotherapy by polymeric micelles and/or nanoemulsion/microbubble encapsulated drugs may offer an innovative approach to PDA treatment. Important physical properties of PFP nanoemulsions allow drug encapsulation, tumor-targeting, enhanced intracellular drug delivery, and enhanced tumor visibility. The success of our pilot studies4 suggests that drug delivery in PFP nanoemulsions combined with non-invasive or minimally invasive ultrasound-mediated chemotherapy may allow development of curative rather than palliative therapy for pancreatic cancer. Here we present and discuss advantages and limitations of this approach for therapy and imaging of human MiaPaCa-2 pancreatic cancer xenografts inoculated in nu/nu mice. Cancer cells were transfected with red fluorescent protein (RFP) to allow comparison of RFP and ultrasound imaging data.
Block copolymers used in this study were bought from Polymer Source Inc. (Montreal, Quebec, Canada). The PEG-PLLA copolymer had a total molecular weight of 9,700; the molecular weights of a hydrophilic PEG block and a hydrophobic PLLA block were 5,000 D and 4,700 D respectively, the number of the monomer units in the corresponding blocks was 113.6 and 54.7.
Micellar solutions of the PEG-PLLA and PEG-PCL block copolymers were prepared by a solvent exchange technique as described in detail previously2. Genexol-PM (GEN) was obtained from Samyang Corp. (Daejeon, South Korea). Genexol PM is a powder formulation that contains 100 mg PTX in 850 mg of Genexol PM powder (other components are mPEG-PLA and lactose). A desired weight of the GEN powder was dissolved either in PBS or in the PEG-PLLA micellar solution; this composition is designated mGEN.
Paclitaxel-loaded PFP nanoemulsions (nbGEN) were prepared as follows; 1% vol. PFP was added to mGEN and samples were sonicated in ice-cold water by 20-kHz ultrasound until all PFP was transferred into an emulsion. GEM was added to this composition in a desired concentration.
Size distribution of nanoparticles was measured by dynamic light scattering at a scattering angle of 165° using Delsa Nano S instrument (Beckman Coulter, Osaka, Japan) equipped with a 658 nm laser and a temperature controller. Particle size distribution was analyzed using Non-Negative Least Squares (NNLS) method.
Unfocused 1-MHz ultrasound was generated by an Omnisound 3000 instrument (Accelerated Care Plus Inc, Sparks, NV). Pressure generated by the transducer was measured with ONDA needle hydrophone (Onda HNR-0500). Mechanical index (MI) was calculated according to the equation MI=Pr.3/F0.5, where peak rarefactional pressure Pr = 0.61 MPa and F = 1 MHz. In our experiments, MI = 0.59 was used.
Pancreatic cancer MiaPaCa-2 cells were obtained from American Type Culture Collection (Manassas, VA). Cells were maintained in DMEM supplemented with 10% heat-inactivated fetal bovine serum (FBS). MiaPaCa-2 cells were transfected with red fluorescence protein (RFP) according to the procedure described in refs.56, 57. Cells were cultured at 37 °C in humidified air containing 5% CO2.
Orthotopic pancreatic cancer was inoculated surgically in the pancreatic tail57. Mice received a single sub-capsular injection of 106 red fluorescent protein labeled MiaPaCa-2 cells suspended in 0.125mL serum free media (DMEM). All procedures were done utilizing a 12X Universal S3B microscope (Carl Zeiss, Thornwood, NY).
Following the primary surgery, high resolution (3456 pixels X 2304 pixels) whole body digital images (EOS Digital Rebel, Canon USA, Lake Success, NY) of each mouse were obtained once a week to monitor primary tumor growth and presence of metastases. The red fluorescent protein was visualized with an Illumatool Bright Light System that consisted of a 563nm excitation filter and a 587nm emission filter (Model LT-9900, LightTools Research, Encinitas, CA). Animals were imaged under nose-cone induced isoflurane general anesthesia. Primary tumor area was quantified using public domain software ImageJ (National Institutes of Health http://rsb.info.nih.gov/ij/).
Animals were randomly assigned to six groups, six animals each: (1) negative control (injection of PBS); (2) GEM at 140 mg/kg (positive control 1); (3) GEN at 20 mg/kg as PTX (positive control 2); (4) GEM+GEN combination treatment; (5) GEN + ultrasound; (6) GEM + nbGEN + ultrasound. Unfocused continuous wave 1-MHz ultrasound was applied extracorporeally for 30 s to the pancreas region of abdomen. Nominal ultrasound intensity was 3.4 W/cm2, which corresponded to a measured MI = 0.59. Ultrasound was applied extracorporeally to abdominal area in the pancreas region through a water bag coupled from both sides to ultrasound transducer and mouse skin by the ultrasound coupling gel. Ultrasound was applied 4 to 5 hours after drug injection. In the first treatment round, drug was injected twice a week for two weeks then there was a break for two weeks and the treatment was repeated using the same protocol as in the first treatment round. The dose of PTX was the same in all formulation used in this study.
Ultrasound imaging was performed using a 14-MHz linear transducer (Acuson Sequoia 512, Siemens, Mountain View, CA).
The size distribution parameters for PEG-PLLA micelles and nanodroplets at room temperature are presented in Supporting Material, Table 1 (adapted from ref.4). The size distribution of micelles was bimodal, with smaller particles corresponding presumably to individual spherical micelles while larger particles represent either wormlike micelles or micellar aggregates. For both types of particles, paclitaxel-loaded micelles (29.3 nm) were larger than empty micelles (22.2 nm). The formation of nanoemulsion after the perfluoropentane (PFP) introduction resulted in the disappearance of small micelles and generation of nanodroplets (592.6 nm, 73%). For paclitaxel-loaded systems, introduction of PFP resulted in a tri-modal size distribution. The size of micelles dropped from 29.3 nm to 19.3 nm while the size of nanodroplets increased from 592.6 to 718.4 nm, suggesting paclitaxel transfer from micelles to nanodroplets. Based on this information, in the formulation used in vivo, the entirety of the drug may be considered located in nanodroplets.
Due to a large difference between acoustic impedance of PFP nanodroplets or microbubbles and water or biological tissues, both PFP nanodroplets and microbubbles manifest echogenic properties (i.e. ultrasound backscattering) thus developing ultrasound contrast (Figure 1A–C). However, contrast enhancement by microbubbles is much stronger than that by nanodroplets (compare brightness of bubbles indicated by horizontal arrow in Figure 1C with that of droplets indicated by vertical arrow in Figure 1C). This feature allows discrimination between droplets and bubbles after in vivo injections. When PFP nanoemulsions are injected directly into the tumor, a very strong and long-lasting ultrasound contrast is generated due to a partial droplet-to-bubble conversion and bubble coalescence (Figure 1D)2, 43. The grayscale intensity of tumors also increases after systemic injections of nanoemulsions, indicating nanodroplet extravasation into tumor tissue4, though this effect is not as strong and punctate as after direct intratumoral injections (Figure 2, compare gray scale before (A) and after injection of paclitaxel (PAX)-loaded PFP/PEG-PLLA nanoemulsion (B).
Tumor sonication after nanodroplet accumulation results in an increase in brightness of bright specks that have been already present in the selected tumor slice; formation of new bright specks is visible in their vicinity (compare the sites indicated by arrows in Figure 3 A, B). This suggests ultrasound-induced, highly localized droplet-to-bubble conversion in tumor tissue in the vicinity of pre-existing microbubbles, as was earlier observed in model experiments on microbubbles inserted in the agarose gel or plasma clot4, 43. Though it is difficult to image exactly the same plane at different days, it appears that bright specks are gradually spreading into the lines with slightly increased brightness (Figure 3C). The data presented above suggest that drug-loaded PFP nanodroplets accumulate in the pancreatic tumor tissue after systemic injection and can be converted into microbubbles under the ultrasound action, which is the foundation of the chemotherapeutic modality presented below.
Tumor growth curves for animals treated by Gemcitabine (GEM), micellar formulation of paclitaxel Genexol PM (GEN), combination drug Genexol PM+GEM, and nanoemulsion formulation of paclitaxel nbGEN+GEM and ultrasound are presented in Figure 4. Good correlation was observed between tumor sizes measured by RFP and ultrasound imaging. GEM-treated tumors grew with the same rate as control while those treated with PTX-containing formulations dramatically regressed indicating sensitivity to PTX.
As shown in Figure 4, systemic chemotherapy by nanodroplet encapsulated paclitaxel combined with GEM and ultrasound resulted in dramatic tumor regression. The gemcitabine that was added in this formulation has a very high aqueous solubility, is not internalized by either micelles or droplets and circulates independently.
The effect of the combined treatment with nanodroplet encapsulated Genexol PM, GEM and ultrasound was the strongest among the six treatment protocols reported in the experimental section (for instance, P << 0.001for GEM+nbGEN+ultrasound vs. Genexol PM treatment in a paired T-test) (Figure 4). The ultrasound effect on the treatment by the nanodroplet encapsulated paclitaxel (nbGEN) was stronger than that on micellar encapsulated paclitaxel (Genexol PM)4. Even a very large initial tumor effectively regressed under the combined action of nanodroplet encapsulated paclitaxel and ultrasound (Figure 5). Interestingly, the treatments that involved tumor sonication resulted in a significantly reduced number of metastatic foci and suppression of ascitis formation (Figure 6). Ascitis was clearly visible in ultrasound images of control or GEM-treated tumors (Figure 7); no ascitis was found in images of ultrasound-treated tumors or postmortem. This important effect was unexpected and counterintuitive. Possible reasons of metastasis suppression are discussed below.
Note that for all treatment groups, treatment interruption for two weeks after the first treatment round resulted in tumor re-growth. A second treatment with the same regimen was less effective than the first one4. In the Genexol PM and GEM+Genexol PM groups, residual tumors stabilized but did not regress during the second treatment round; only nanodroplet/ultrasound therapy resulted in some regression of re-grown tumors during the second treatment round. This data suggest that either some resistance to paclitaxel developed in the course of the initial treatment or resistant cells were selected during the first treatment round.
With any treatment protocol, local tumor recurrence was observed after completion of treatment. The local recurrence occurred even when the residual tumor could not be resolved by RFP imaging. The possible reasons of this effect are discussed below.
It is noteworthy that the intra-group variations were larger for micellar or nanodroplet encapsulated paclitaxel groups than for control or GEM group. An example for GEM and nbGEN+GEM+US group is shown in Figure 8.
Experimental results are briefly summarized below.
The tumor growth curve for GEM was very close to that for a control group, confirming GEM-resistance of MiaPaCa-2 cells reported earlier in ref.58. In this work, the GEM resistance of MiaPaCa-2 cells was ascribed to the action of a multidrug resistance associated protein MRP. MRPs are a subfamily of the ATP-binding Cassette (ABC) superfamily of transporter proteins. MRPs are separate and distinct from the well-known family of transporter proteins called P-glycoproteins (P-gp). MRPs are usually localized in plasma membranes59, 60 but may be relocated to the intracellular membranes of late endosomes and lysosomes61. As indicated by its name, the MRP action is expressed against a plethora of drugs. It was shown in ref.58 that GEM-resistant MiaPaCa-2 cells were sensitive to paclitaxel. This was confirmed in the present study. Animal treatment with micellar encapsulated paclitaxel, Genexol PM resulted in a substantial tumor regression (Figure 4). Cell culture experiments showed that MiaPaCa-2 cells were sensitive to dissolved paclitaxel, Taxol®, as well as micellar paclitaxel Genexol PM (data not shown). These data indicate that GEM-resistance of MiaPaCa-2 cells is not a multidrug resistance and is not caused by the action of plasma membrane pumps that would have acted on both paclitaxel and GEM.
GEM and paclitaxel have profoundly different mechanisms of action. GEM is the nucleoside analogue and its site of action is in cell nuclei. On the contrary, paclitaxel acts by stabilizing microtubules in cytoplasm thus mechanically preventing cell division. The sensitivity of GEM-resistant MiaPaCa-2 cells to Genexol PM suggests that paclitaxel loaded in PEG-PLLA micelles successfully overcomes plasma membrane barriers thus allowing paclitaxel interaction with microtubules. GEM-resistance of MiaPaCa-2 cells may be caused by the action of nuclear pumps that exert no effect on paclitaxel. Other possible mechanisms include several genetic and/or epigenetic alterations; the latter include gene products associated with gemcitabine transport and metabolism (for a recent review, see ref. 62).
Low effecacy of GEM has warranted studies of combination drugs63–65. Therefore in this study, we included GEM as a component of a combination formulation with Genexol PM. It was found that combination treatment by GEM+Genexol PM was slightly more effective than Genexol PM alone though the main effect was undoubtedly exerted by Genexol PM (in a paired T-Test, statistically significant differences were manifested between these groups, P = 0.01) (Figure 4). This suggests that Genexol PM may affect intracellular mechanisms involved in GEM inactivation, which deserves further exploration.
The most efficient tumor regression was observed during systemic treatment with nanodroplet encapsulated paclitaxel combined with tumor-directed ultrasound. As suggested by ultrasound imaging, nanodroplets with encapsulated paclitaxel accumulated in tumor tissue. They converted into microbubbles and released their drug load under the action of tumor-directed ultrasound, which resulted in efficient chemotherapy of pancreatic cancer. However, local tumor recurrence was observed during the treatment break or after the completion of treatment and the recurrent tumors proved more resistant to the same treatment regimen indicating developed drug resistance or selection for the resistant cells during the initial treatment. A possible reason for this adverse effect is discussed below.
Ultrasound imaging manifested a highly non-uniform distribution of nanodroplets throughout the tumor volume (Figure 9). This may be caused by the irregularity in tumor vascularization and distribution of inter-endothelial gaps, which may results in intra-group variability. Doppler images showed the blood vessels that could be resolved by the instrument (hundreds of micron size) localized at the tumor periphery or around the tumor (Figure 10); smaller capillaries could not be resolved but tumors are known to have irregular vascularization7, 8.
After drug release, irregularity of nanoparticle extravasation would result in drug gradients within the tumor volume. If this is the case, some tumor sites may be exposed to sub-therapeutic concentrations of drug, which would favor development of drug resistance. This problem may be solved, at least partly, by tumor sonication. Ultrasound is known to enhance diffusion66, 67. In our earlier work with micellar doxorubicin we found that the intratumoral drug distribution was significantly more uniform in sonicated tumors6. However, the degree of drug equalization would depend on the mechanical properties of tumor tissue and tumor vascularization. The above problem may be pertinent to any nanoparticle-associated drug delivery modality. In order to suppress or prevent the development of drug resistance, introduction of MDR-suppressing agents into nanoparticle drug formulations may be warranted.
Another intriguing effect that deserves exploration is related to the suppression of pancreatic tumor metastases by ultrasound-mediated chemotherapy with micellar- or nanoemulsion encapsulated paclitaxel (Figure 6). Recent works have revealed that mechanical forces can profoundly influence cell behavior by affecting cell spread, growth, survival and motility. For instance, it was shown that tumors in the pancreas display compromised laminin and type IV collagen basement membrane organization and thinning that, when combined with outward projecting compression force, facilitates tumor cell invasion into the parenchyma68. Research on the mechanical components of tumor growth, survival and motility is still in its infancy68–70. A better understanding of how mechanical force signals regulate tumor progression and metastasis could significantly accelerate the development of improved treatments.
Transmembrane integrins are the most likely mechano transducers. Two families of proteins are currently implicated in this function: Rho (a member of the Ras protein superfamily)71, 72 and PINCH73. The latter functions as an adapter at a key convergence point for integrin and growth factor signal transduction. It was shown that PINCH is up-regulated in the stroma of common cancers, notably at invasive edges. This seems important, as the dense desmoplastic infiltration within the tumor is a unique characteristic of PDA. Since PINCH has been shown to play a role in cell migration and signal transduction, PINCH expression may serve a potential marker for tumor invasion. Research on animal models and pancreatic human tissue revealed that PINCH expression is associated with a more aggressive pancreatic tumor and shorter survival57. We hypothesize that reduction of metastases in animals treated by Genexol PM or nbGEN and ultrasound may be associated with a concomitant reduction of PINCH expression due to the ultrasound effect on cell membranes. These experiments are currently in progress.
In conclusion, for effective PDA therapy, paclitaxel delivery in PFP nanoemulsions may be combined with endoscopic, extracorporeal, or even intraductal ultrasound applicators. Our experiments demonstrated that low-power output application of ultrasound was able to release the drug into tumor tissue2–4. Experimental intraductal transducers are being developed in France; these would allow the most precise tumor sonication. A substantial reduction of the required ultrasonic energy would solve a number of difficult technological problems for the intraductal application. Recent advances in focusing extracorporeal ultrasound on the pancreas74 suggests that this approach to pancreatic sonication may also become available clinically in the near future.
Ultrasound-mediated therapy with the combination drug formulation GEM + nanodroplet-encapsulated paclitaxel nbGEN combined with tumor sonication with 1-MHz ultrasound exerted strong therapeutic effect and produced dramatic regression of even very large PDA tumors. Ultrasound-mediated therapy resulted in suppression of metastases and ascites suggesting improved control of invasive cancer cells. Tumor relapse was observed after the completion of therapy and recurrent tumors proved more resistant to the same therapy indicating development of resistance, which may be caused by drug gradients in tumor volume and presence of zones with sub-therapeutic drug concentration. This problem may be pertinent to any therapy that relies on nanoparticle extravasation and may require the addition of MDR suppressing agents into nanoparticle drug formulations.
Table 1 of Supporting Material presents particle size distribution parameters for empty and PTX-loaded PEG-PLLA micelles and nanodroplets.
AVI Video file of Supporting Material shows scanned ultrasound images of a pancreatic tumor taken 5 hours after the systemic injection of PTX-loaded PEG-PLLA nanodroplets.
The work was supported by Grants Number R56EB001033 and R01EB001033 to NR from the National Institute of Biomedical Imaging And Bioengineering. The content is solely the responsibility of the authors and does not necessarily represent the official views of the National Institute of Biomedical Imaging And Bioengineering or the National Institutes of Health.
The information is available free of charge via the Internet at http://pubs.acs.org