(718e) Controlled Release of Drug-Loaded Microbubbles in Blood Vessels On a Chip

Ultrasound contrast imaging using micron-sized bubbles (less than 8 µm in diameter) has demonstrated detectable echogenic signals in vivo. However, challenges remain in terms of producing monodisperse microbubbles stable against destruction from aggregation and gas dissolution for improving blood circulation times and reducing clogging of small blood vessels. In the case of drug delivery, it is crucial to control destruction to release encapsulated contents only to the targeted area with minimal prior passive leakage of drug. By using a photopolymerizable diacetylene lipid as a component in a microbubble shell matrix, the surface properties such as elasticity or resistivity against gas dissolution can be tuned,¹ which therefore enables imaging at one frequency and drug release in the area of interest at a different frequency. The bubbles will be transported into in vitro blood vessel structures formed on a microfluidic device to examine the effect on cells under ultrasound. Polymerizable lipid mixtures, consisting of ethylene glycol diacetylene lipids (h-PEG₁PCDA), PEG-diacetylene lipids (m-PEG-PCDA) with PEG 5000 (5K) and L-α-phosphatidylcholine, hydro soy PC were prepared by varying  the amount of h-PEG₁PCDA from 0 to 15 mol% and by keeping m-PEG-PCDA constant at 15 mol%, with the balance consisting of hydro soy PC. Oil-layered drug-encapsulated microbubbles were produced using a double-emulsion microfluidic focusing technique. Anti-cancer drug doxorubicin (DOX) and paclitaxel (PAX) were used. For liposomal-drug tethered microbubbles, the polymerizable lipid was also used. The in vitro blood vessel model on a chip was adapted from Kim et al.² The monodisperse Polymerized Shell Microbubbles (PSM) (d_avg = 2.2 µm, polydispersity = 8%) containing 25 mol% of polymerizable diacetylene lipids (25%DA: 10 mol% h-PEG₁PCDA and 15 mol% m-PEG₂₀₀₀PCDA) were more stable than commercially acquired microbubbles (Vevo MicroMarker, VMM) or non-polymerizable shell microbubbles (NSM) in terms of dissolution. The 25%DA showed a significantly slower decrease in intensity of gray-scale ultrasound image brightness than VMM or NSM. The half-life (t_H the time for the ultrasound signal to drop 6 dB in brightness) for the 25%DA was t_H = 5 min, and for both VMM and NSM, t_H < 1 min (data not shown). In addition, the bubbles that were polymerized to different extents (t_polym) showed variable destruction rates at different ultrasound power levels, suggesting that polymerization can not only provide passive bubble longevity but tunable rupture capability. Various methods of drug loading on the microbubbles were demonstrated. The double-emulsion microfluidic device produced oil-layered PAX-encapsulated microbubbles. Drugs encapsulated in liposomes formed using a polymerizable lipid shell showed drug-release upon ultrasound and further demonstrates the effect of the polymerized shell on shell resistance against drug leakage. The polymerized liposomes showed about 0% drug released compared to the unpolymerized liposomes with 21% of total drug encapsulated in the liposomes at 3.5 MHz for 1 min insonation. Lastly, the microbubbles in the in vitro blood vessel on a chip were observed under ultrasound application. The size distribution of the microbubbles in the device increased as the insonation time increased, indicating the in vitro microfluidic device provides a good model to examine the interaction between the drug-encapsulated microbubbles and the cells under ultrasound. These results indicate that the dissolution of microbubbles in the bloodstream or under ultrasound stimulation is tunable by varying the fraction of polymerizable lipid or polymerization time. Therefore, these tunable microbubbles have the potential to be customized ultrasound contrast agents for targeted molecular imaging and therapeutic treatment applications.

[1] Park et al. Langmuir 2012, 28, 3766-3772. [2] Kim et al.  Lab Chip 2013, 13, 1489.