Light-Activatable Theranostic Agents for Image-Monitored Controlled

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Light-Activatable Theranostic Agents for Image-Monitored Controlled Drug Delivery Zhe Zhang, Madison Taylor, Courtney Collins, Sara Haworth, ZhanQuan Shi, Zheng Yuan, Xingyu He, Zishu Cao, and Yoonjee C Park ACS Appl. Mater. Interfaces, Just Accepted Manuscript • DOI: 10.1021/acsami.7b15325 • Publication Date (Web): 25 Dec 2017 Downloaded from http://pubs.acs.org on December 25, 2017

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ACS Applied Materials & Interfaces

Light-Activatable Theranostic Agents for Image-Monitored Controlled Drug Delivery Zhe Zhang1, Madison Taylor1, Courtney Collins1, Sara Haworth1, ZhanQuan Shi2, Zheng Yuan1, Xingyu He1, Zishu Cao1, Yoonjee C. Park1* 1

Department of Biomedical, Chemical & Environmental Engineering, 2College of Pharmacy, University of Cincinnati, Cincinnati, OH, USA *Corresponding Author: [email protected]

Abstract：： A novel drug delivery vehicle using nanodroplets activated by light irradiation for drug release in a controlled manner has been developed. Drug encapsulated in the nanodroplets was released upon the phase-transition from liquid droplet-to-microbubbles (vaporization) by plasmonic photothermal heat from gold nanorods adsorbed on the surface of the nanodroplets. The nanodroplets were stable against aggregation and dissolution at 4 ºC over 3 months to date. The phase-transition was quantitatively analyzed by ultrasound imaging to examine the amount of drug release non-invasively. In vitro studies showed that cell death occurred only when light irradiation was performed with the drug-encapsulated nanodroplets. Ex vivo studies demonstrated a potential application of the nanodroplets for treating posterior eye diseases. Thus, it has been demonstrated that our gold nanorod-coated light-activatable nanodroplets can be a candidate for a controlled release and a dosage-monitored drug delivery system.

Keywords: Light-Activatable, On-demand Drug delivery, Ultrasound Imaging, Posterior Eye Diseases, Plasmonic Photothermal Effect.

1. Introduction Drug delivery systems for temporal and spatial drug release using nano-carriers, or so called “on-demand” drug delivery systems, are a new therapeutic approach in medicinal fields. Drug release can be tightly controlled and negative side effects can be minimized. For instance, such therapies can be applied locally for wound treatment via controlled release of anti-inflammatory drugs and/or growth factors from nanoparticles to selectively induce tissue formation 1-2. Active release via external triggering devices is useful to control when and how much drug is delivered at a target of interest. Ultrasound, especially high-intensity focused ultrasound (HIFU), has been widely investigated to control release of drug utilizing phase-shift nanodroplets, called acoustic droplet vaporization 3-5. Laser or light can be another promising triggering device because they allow more precise control of drug release and targeting lesions with minimized adverse effects to surrounding tissue by choosing duration, energy, or beam area, compared to other methods such as HIFU or local heating 6-7. In addition, laser can be used for gaseous enclosures in the human body such as the lungs and pressure-sensitive organs such as the eyes, whereas ultrasound beams cannot effectively be used 8. Recent studies have shown combination of the droplet phase-transition with light, utilizing photothermal nanoparticles, called optical droplet

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vaporization, for biomedical applications 8-9. One study showed feasibility of delivering drug upon laser irradiation in vivo using phase-transition nanoparticles 10. However, wavelengths of visible light (400 – 700 nm) have been mostly used, which cannot penetrate deep into tissue and often damage susceptible tissues in the skin or the eyes. In this study, we have developed nanodroplets which are responsive to a near-infrared laser, capable of sufficiently penetrating tissue, to reach target lesions in the human body 11. We have used gold nanorods with a surface plasmon resonance at the near-infrared region (700 – 2500 nm). Plasmon resonance nanoparticles are known to generate heat after photon absorption 12. Our gold nanorod-coated nanodroplets release drug upon vaporization of the liquid core which then undergo phase-transition to bubbles via laser irradiation because of heat generated on the nanodroplet surface (Figure 1). The details on each component and synthesis process are in Section 2.1. Perfluorocarbon gas bubbles, which can be obtained from phasetransition of the nanodroplets, are clinically used as ultrasound contrast agents. Thus, the phasetransitioned gas bubbles allow us to monitor the drug release via ultrasound imaging. 13 Pulsed irradiation of light to gold nanoparticles compared to continuous wave (CW) provides various advantageous outcomes, including nanoscale control of temperature distribution 14, photothermal cancer therapy 15, and drug release 16. For example, a femtosecond laser can trigger a sudden temperature increase at the sub-nanosecond scale by a high short peak energy 17. This contributes to confining the temperature increase close to the gold nanorods to avoid overheating of the surround medium 17-18. Despite these advantages, pulsed lasers may cause cumulative temperature increase greater than needed, which can lead to undesirable tissue damage. In this study, we have measured temperature change after pulsed laser irradiation at different laser parameters and optimized to avoid overheating. When the nanodroplets undergo phase-transition (vaporization) to gas bubbles, the bubbles dramatically enhance ultrasound imaging contrast due to their different impedance and resonance with ultrasound waves, called echogenicity 19. In our developed system, the amount of drug released from nanodroplets was approximately calculated by correlating the brightness of the ultrasound images. As a result, delivery of drug can be monitored and confirmed qualitatively and quantitatively. In this study, we have demonstrated feasibility of applying our light-induced drug delivery system to posterior eye disease treatment, such as macular edema, retinopathy or retinoblastoma. Overall, we have developed and investigated nanodroplets that are potential “on-demand” theranostic agents for imagemonitored controlled drug delivery.


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Figure 1. Schematics of gold nanorod-coated phase-transition nanodroplets.

2. Materials and Methods 2.1.

Nanodroplet Synthesis, Materials, and Preparation

The nanodroplets were synthesized via a double emulsion process. The first emulsion of an anti-cancer drug, melphalan (Tocris Bioscience), aqueous solution dispersed in the liquid perfluorcarbon, C5F12 (perfluoropentane) (Synquest Laboratories) was prepared with Krytox (DuPont) as polymer surfactant via probe sonication in an ice bath. The second emulsion was then created by combining citrate-coated gold nanorods (Nanopartz) and liposomes to the first emulsion and sonicating in an ice bath. The TEM image of gold nanorods is in Figure S1(A) Supporting Info. The resulting solution was washed three times by centrifugation (2700 rpm) and replacing supernatant with water. The gold nanorods’ dimensions were 10 nm in diameter and 59 nm in length and the concentration used was 4.19E+11 gold nanorods per mL. The liposome components, consisting of stearylamine (Tokyo Chemical Industry Co., Portland, OR), 1,2distearoyl-sn-glycero-3-phosphoethanolamine-N-[methoxy(polyethylene glycol)-5000] (DSPE-PEG 5K) (Nanocs, Boston, MA), and 1,2-dipalmitoyl-sn-glycero-3-phosphocholine (DPPC) (Avanti Polar Lipids, Inc., Alabaster, AL) at 50 mol %, 15 mol%, and 35 mol %, respectively, were dissolved in chloroform (10 mg/ml). The mixture was evaporated in a hood and the dried film was hydrated with distilled water and sonicated at room temperature for 10 minutes to form liposomes. For the targeted nanodroplets 1,2-distearoyl-sn-glycero-3-phosphoethanolamine-N-[amino(polyethylene glycol)-5000 (DSPE-PEG 5K-NH2) was used instead of DSPE-PEG 5K in the liposome formulation described above. Folic acid (Sigma Aldrich) was functionalized with N-(3-dimethylaminopropyl)-N’ethylcarbodiimide hydrochloride (EDC) and N-hydroxysuccinimide (NHS), in pH 9 overnight on an agitator at room temperature. The folate-functionalized nanodroplets were obtained after dialysis for 48 h against deionized (DI) water at pH 9. Thus, the polymer linker depicted in Schematics of Figure 1 is PEG 5K.



2.2. Characterization – Optical Imaging, Spectroscopy, and Zeta Potential The size, drug encapsulation, and gold nanorod coating were determined by optical microscopy and spectroscopy. Differential interference contrast (DIC) and fluorescence microscopy (Nikon, Inc.) were used to determine the size and drug/dye encapsulation of the nanodroplets optically. Dynamic light scattering (NanoBrook Omni, Brookhaven Instruments Cooperation) was used to determine the hydrodynamic diameter of the nanodroplets. A UV-vis spectrometer (SpectraMax, Molecular Devices, LLC) and a Raman microscope (inVia, Renishaw, plc) were used to detect gold nanorod coating and drug or dye encapsulation. An enhanced darkfield hyperspectral microscope (CytoViva, Inc.) was used to confirm gold nanorod coating on individual nanodroplets. Zeta-potential was measured to confirm gold nanorod adsorption on the nanodroplet surface (NanoBrook Omni, Brookhaven Instruments Cooperation).

2.3. Stability of Nanodroplets Dynamic light scattering (DLS) at 25°C was used to monitor stability of the nanodroplets against aggregation or rupture by size measurement. 150 µl of the nanodroplets in water solution was diluted in 1.35 ml of DI water for the measurements. Three measurements were taken for each sample and averaged. The amount of nanodroplets left was quantitatively determined by light absorbance of the nanodroplets at 633 nm. With different dilution factors, the measurement values were linearly proportional to the concentrations. To determine stability against drug leakage to the surrounding medium from the nanodroplets, melphalanencapsulated nanodroplets, empty nanodroplets (nanodroplets which do not contain melphalan), water, and gold nanorod solutions were stored in a water bath at 37 ºC. The drug (melphalan) concentration in the supernatant after centrifugation at 10,000 g force was measured using UV-VIS spectrometry (SpectraMax, Molecular Devices) at 261 nm peak position of the melphalan in DI water.

2.4. Laser Activation of Nanodroplets (Phase-transition) A picosecond pulsed laser (WedgeHF, RPMC Lasers Inc.) at 1064 nm, 1.8 W average power, 500 ps, and 10 kHz pulse repetition frequency was used throughout the phase-transition and drug release experiments except the phase-transition observation experiment with an optical microscope. The density of the perfluorocarbon is 1.63 g/cm3 20; thus, the nanodroplets settled down within an hour by gravity. For the optical observation experiment, a Mai Tai (Spectra Physics) femtosecond laser at 980 nm, 0.08 W average power (measured), 100 fs, and 80MHz pulse repetition frequency was used.

2.5. Ultrasound Imaging A Terason 3200t ultrasound imaging system (Terason, Burlington, MA) was used for monitoring the phase-transition in real-time. The transducer frequency and the mechanical index used were 16 MHz and 0.1, respectively. A sample in a 1.5 mL test tube was immersed in a tank filled with DI and degassed water and positioned with the picosecond pulsed laser. The transducer was positioned at a 90 degree angle to the laser. The ultrasound imaging was recorded using software Snagit (TechSmith). Screen shots were taken to quantitatively analyze the brightness change of a selected region of interest (ROI) over time using the software ImageJ (National Institutes of Health). By using the plotting Z-axis profile function of 4 ACS Paragon Plus Environment

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ImageJ, the profile of brightness of the ROI was obtained. To quantitatively analyze the plot, the x-axis values and y-axis values were input into Excel, and we adopted a trapezoidal approximation method to calculate the area of the peak, which corresponds to the amount of drug released in phase-transition.

2.6. Ex Vivo Phase-Transition and Ultrasound Imaging (Pig Eyes) Ex vivo phase-transition tests were conducted in pig eyes (Carolina Biological Supply Company). Before conducting the experiments, we cut some skin and fat off the eyeball using a razor blade to show the lens and to make the outer surface smooth for the convenience of observation and fixation. We examined the eyeball before injection using the ultrasound imaging system with the lens facing upward and marked the position for consistent measurement for each trial. After injection of the 100 µl nanodroplet dispersion, the eyeball was examined the same way by positioning the ultrasound transducer at the same position. Then, the laser was irradiated vertically through the lens multiple times of 20 s irradiation to irradiate all the nanodroplets injected. Afterwards, the eyeball was examined again using the ultrasound imaging system (n=3). Distribution of the nanodroplets that encapsulate dye (CF680R, Biotium) upon irradiation was detected by fluorescence measurement in the vitreous humor via tissue dissection. After irradiation, the eyeball was stored in a -70 °C freezer for 40 minutes with the lens facing upwards. Then the eyeball was cut into 8 segments mutually perpendicular (Schematics in Figure 6A) and stored in 8 different numbered glass vials for the vitreous humor to melt. Finally, fluorescence was measured for each segment. The data shown in Figure 6B were obtained after background subtraction using the fluorescence value of the vitreous humor as a background. The effect of freeze-thaw cycle on the nanodroplet stability against rupture was also examined (Supporting Info for the details, Text S1).

2.7. Drug/Dye Release Analysis Drug release upon light activation was quantified by measuring fluorescence of drug (doxorubicin) in the nanodroplets after dissolving them in acetonitrile using a plate reader (SpectraMax, Molecular Devices). After light irradiation, supernatants of the nanodroplet dispersions were removed and the remaining particle sediments were dissolved in acetonitrile. The percentage release was calculated based on the values. Supernatants of the nanodroplet dispersions were collected for measurements before and after light irradiation. A calibration curve of doxorubicin fluorescence was generated.

2.8. Encapsulation Efficiency of Drug Encapsulation efficiency of drug was calculated based on the following formula:

% =

100 ℎ

Theoretical drug concentrations were calculated based on volume fraction of drug solution used and the initial drug concentration. Drug concentrations were determined by UV-Vis absorbance and their calibration curves. Calibration curves for doxorubicin (Dox) and melphalan were obtained by UV-Vis absorbance in a 1:1 DI water:acetonitrile solution. Concentration of the drug in nanodroplets was



determined by UV-Vis absorbance after dissolving the nanodroplets dispersed in DI water in the same volume of acetonitrile.

2.9. In Vitro Drug Delivery Efficacy Human Retinal Pigment Epithelium (RPE) cells (ARPE-19, ATCC) were cultured in Dulbecco's Modified Eagle Medium: Nutrient Mixture F-12 (DMEM/F-12) supplemented with 10% fetal bovine serum (FBS) and 1% Penicillin-Streptomycin Solution at 37° C, 5% carbon dioxide. RPE cells were seeded in 96 well plates with a density of 10,000 cells in 100 µl medium per well. When the cells were confluent (24 hrs), 100 µl volume of sample dispersed in medium was added to each well and the picosecond pulsed laser was irradiated directly to the cells. After 24 hr incubation, XTT cell proliferation assay (ATCC) was added to measure the cell viability. The free melphalan concentration used in the in vitro test was 75 µg/mL and determined based on our previous data of cytotoxicity at different concentrations of melphalan to RPE cells (Supporting Info, Figure S2). 50~300 µg/mL have been used to treat retinoblastoma in human studies 21-22.

2.10. Temperature Measurement After Laser Irradiation Temperature of the nanodroplet dispersion after laser irradiation was measured by two different methods: 1. using a temperature probe, and 2. using an infrared (IR) camera. For the temperature probe method, a 1:10 dilution of the nanodroplet dispersion in DI water (40 µL) was prepared in a micro centrifuge tube. The tube was clamped approximated 3 inches in front of the laser. The nanodroplet dispersion was then irradiated by the laser for a designated amount of time (20s, 40s, etc.). Once the laser was powered off, the temperature was measured (n=3). For the IR camera observation method, 1 ml of the nanodroplet dispersion in a 1.5 ml test tube was irradiated by the laser. An infrared camera (T640, FLIR) was positioned at a 90 degree angle to record temperature change during the irradiation. The temperature was read based on the color bar on the camera screen.

3.

Results


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3.1.

Nanodroplet Characterization

Figure 2. (A) Overlay of fluorescence and DIC (Differential Interference Contrast) optical images (60x); (B) (a) UV-vis absorption (b) Raman spectrum. Blue line: gold nanorod solution, red line: gold nanorod-coated nanodroplets, green line: the single emulsion (water droplets in perfluorocarbon); (C) Enhanced dark field images; (D) Hyperspectral image analysis of the nanodroplets and nanodroplets with no gold.

The average size of the nanodroplets measured by optical images and by dynamic light scattering was ~250 nm and ~300 nm in diameter, respectively (Figure 2A and 3, respectively). When Rhodamine B dye was co-encapsulated with drug, the encapsulation inside the nanodroplets was visualized by fluorescence images (Figure 2A, red color in the nanodroplets). UV absorption scanning also proved presence of the dye in the nanodroplets as dissolved in an aqueous phase (W1 in W1/O/W2 emulsion) by showing the peak at a wavelength of 555 nm (Figure 2B (a)) 23. Gold nanorods on the nanodroplets were observed via enhanced darkfield imaging and hyperspectral analysis (Figure 2C and D). The characteristic absorption peak near 510 nm was detected at the surface of individual nanodroplets while bare nanodroplets without gold nanorods on the surface did not show any peak. Raman spectroscopy results of the nanodroplets matched those of the gold nanorods, indicating presence of the gold nanorods on the nanodroplets (Figure 2B (b)). Lastly, the zeta-potential results before and after adding the gold nanorods dispersed in DI water were +15 mV and -20 ~ -30 mV, respectively, which indicates electrostatic adsorption of the gold nanorods to the surface.

3.2.

Stability of Nanodroplets 7 ACS Paragon Plus Environment


The gold nanorod-coated nanodroplets were stable against aggregation and dissolution for at least 100 days at 4°C and 21 days at 37°C in DI water to date based on effective diameter measurement results from dynamic light scattering (Figure 3). The nanodroplets at 37°C gradually showed phase-transition to bubbles, resulting in disappearance of the nanodroplets over time. When the nanodroplets were stored in physiological conditions, in a 150 mM NaCl solution at 37°C, aggregation of the nanodroplets was observed approximately after 15 days. Loss of the nanodroplets due to phase-transition was also observed. However, approximately 50% of the nanodroplets remained over 3 weeks, determined by turbidity of the emulsion (Supporting Info Figure S3). For the stability against passive drug leakage, melphalan concentrations in the supernatant after centrifugation were measured by UV-Vis spectroscopy. The UV absorbance value of melphalan in the supernatant from the melphalan-encapsulated nanodroplets was higher by 0.005 compared to empty nanodroplets (nanodroplets which do not encapsulate drug) over a month. UV absorbance of 0.12 at a wavelength of 261 nm corresponds to 1.7 µg/mL of melphalan. Thus, the drug leakage was not significant at least for a month.

Figure 3. Effective diameters of the nanodroplets over time at 4 ⁰C and 37 ⁰C in DI water.

3.3. Phase-transition of Nanodroplets by Pulsed Laser


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Phase-transition upon laser irradiation was observed optically (Figure 4). The phase-transition from nanodroplets (liquid) to micron-sized bubbles occurred only in the red box area where a pulsed laser scanned once. The phase-transition started to be observed at 0.15 second (150 ms) of irradiation. We note that the time shorter than 0.15 second was not possible due to technical limitation. The volume expanded due to the phase-transition. In diameter, 200~500 nm nanodroplets became 1~5 µm bubbles at 0.15 s. Discussion regarding volume expansion is described in the Discussion section. Over time, the bubbles floated and some bubbles shrunk or merged. When the nanodroplet emulsion was diluted and an individual nanodroplet was scanned by the pulsed laser, it was confirmed that the gas formation occurred only from the nanodroplet, not from the surrounding medium. In addition, gold nanorods only did not show bubbles. These suggest that the bubbles were not created due to overheating of the surrounding medium 24. Also the fluorescence from the dye encapsulated in the nanodroplet after phase-transition was observed in the bubble. The results suggest that the contents are still inside or visually near the surface of

Figure 4. Optical observation of phase-transition of the nanodroplets. The last image shows a micron-sized bubble, phasetransitioned from a nanodroplet. It is an overlay of fluorescence and DIC, showing location of the dye.

the micron-sized bubble until it ruptures. The phase-transition property of the nanodroplets was not affected by “age” of the nanodroplets if they are determined to be colloidally stable by dynamic light scattering. We have not observed noticeable change in phase-transition over time through optical images, ultrasound images, or visual observation of bubble formation (data now shown).

3.4. Ultrasound imaging 9 ACS Paragon Plus Environment


The phase-transition process was monitored in real-time by ultrasound imaging. In the supplementary video (Video S1), the B-mode ultrasound images showed brightness intensity changes in the center of the tube due to the bubble formation upon the laser irradiation 5. The brightness intensity change over time was plotted in the chart next to the video (Supporting Info, Figure S4(B)). The bubble formation was detected when the laser irradiation started at 10 s. During 20 s of irradiation, bubble formation continued. The bubbles lingered for the next 10 s after stopping irradiation and the brightness disappeared at ~40 s. We have noticed that areas under the curve of brightness intensity based on the baseline (definite integrals) are proportional to the irradiation time. The representative plots are shown in Figure 5. The average area from the 20 s irradiation plot was twice larger than the one from the 10 s irradiation plot (215.78 vs. 100.49) (n=3). The results suggest that the amount of phase-transition can be quantified by ultrasound image analysis, and controlled by laser irradiation time. In other words, drug release also can be quantified non-invasively in real-time and be tuned by irradiation duration.

Figure 5. Areas under the brightness intensity curves are proportional to the laser irradiation time. (a) 10 s irradiation; (b) 20 s irradiation.

3.5. Ex-vivo test (pig eyes) The phase-transition occurred in ex-vivo pig eyes upon laser irradiation through the lens. The initial fluorescence value of the nanodroplets injected in the bottom section (segment 5 and 7) of the eyeball (Figure 6A) was approximately 1000. After irradiation, 74, 93, 77, and 55 from the segment 1,2,3, and 4, respectively, was detected in the upper section of the eyeball (Figure 6B). Fluorescence intensity, ~200, was detected at each segment from 5 to 7. The data imply that the nanodroplets were distributed to wider area upon irradiation compared to the initial state. The phase-transition was confirmed by B-mode 10 ACS Paragon Plus Environment

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ultrasound imaging (Figure 6C). Brightness increased significantly after laser irradiation. In addition, the area that the phase-transitioned bubbles covered increased from the initial state due to volume expansion of the nanodroplets. The results suggest the nanodroplet drug delivery system can be applied to posterior eye disease treatment and drug release can be detected non-invasively.

Figure 6.(A) Schematics of ex vivo light triggering and tissue dissection for 8 segments of the eyeball. (B) Fluorescence measurement of different sections of an eyeball after laser irradiation (C) Phase-transition of the nanodroplets in ex-vivo pig eyes monitored by B-mode ultrasound imaging.



3.6. Drug Release and Efficacy For drug release quantification, doxorubicin, another anti-cancer drug, was used because of its native distinctive fluorescence property and an orange-red color with which drug release could be visualized by the naked eye. When the drug-encapsulating nanodroplets in a test tube were irradiated through the air for 20 s, approximately 30% of the initial concentration was released (an average of n=5). The encapsulation efficiency of doxorubicin and melphalan in the nanodroplets was 50% and 11%, respectively. To determine cytotoxicity and drug delivery efficacy of the nanodroplets to the cells, empty nanodroplets, melphalan-encapsulating nanodroplets, and a free melphalan solution with the same concentration of the total nanodroplets applied, were incubated for 24 hr after 20 s laser irradiation (Figure 7). The empty nanodroplets did not cause any cell death, similar to the negative control (Cells in Figure 7). However, the melphalan-encapsulating nanodroplets killed about 15%. This is approximately one third of the free melphalan value, 40%. With 20 s irradiation, a fraction (~1/3) of the nanodroplets underwent phasetransition and delivered drug to the cells. These cell viability results are consistent with the quantitative drug release analysis results.

Figure 7. Cell viability after 24 hr incubation with the nanodroplets with or without laser treatment. “Nanodroplets” refers to empty nanodroplets coated with gold nanorods that do not contain drug. Free melphalan of the same concentration that the total melphalan nanodroplets contain was tested. Cells refer to a negative control without any sample.

3.7. Targeting The nanodroplet surface has been modified to have targeting ability. Because folate receptors are present on RPE cells, folic acid was tethered at the end of PEG linker on the surface of the nanodroplet. Figure 8 shows the targeted nanodroplets (red color) bound specifically on the cells (nuclei: blue color)25-26. On the other hand, the non-targeted nanodroplets did not show any localization on the cells and were randomly distributed. The results suggest that the nanodroplets can be targeted to the area of interest.


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Figure 8. Non-targeted nanodroplets vs. Targeted nanodroplets with folic acid. The red color shows nanodroplets and the blue color shows RPE cell nuclei. The overlay images in the bottom demonstrate localization of the targeted nanodroplets on the cells.

4.

Discussion

Overall, the results from this study demonstrate that stable nanodroplets can deliver drug on demand with quantitative diagnostic capability. The materials used to synthesize the nanodroplets were safe based on the in vitro data. The liquid core used for the nanodroplet synthesis is FDA-approved 27. Gold nanorods have two light absorption peaks due to surface plasmon resonance and the two absorption peaks from the gold nanorods used in this study are 510 nm and 980 nm along the short and the long axes, respectively. The nanorods have a broad peak range near 980 nm, partly due to inhomogeneity of the long axes length. The peak intensity at 1064 nm was approximately 75% (Supporting Info, Figure S1(B)) of the peak intensity at 980 nm. A pulsed laser was used in this study to utilize high laser power condensed in a short time (pulse) as well as to avoid heating surrounding media. Heat dissipates between pulses; thus, heat accumulation in surrounding media is much slower compared to when a continuous wave laser is used 17. We have determined maximal duration of the pulsed irradiation at 10 kHz to keep temperature increase of the surrounding liquid, below 3 ºC, not to exceed 40 ºC, where tissues start to get damaged, from an initial 37 ºC (body temperature). The temperature change was proportional to the laser irradiation duration and 20 s of irradiation yielded about 3 ºC increase (Figure 9). As a result, we have used 20 s duration for drug release analysis and in vitro drug delivery tests to avoid any damage due to heating in Section 3.5. This parameter depends on the average power of a laser, pulse repetition rate, and pulse duration.



We note that the phase-transition process is very local and individual based on the optical observation (Figure 4). When the nanodroplet dispersion was diluted and scanned by a laser, we observed that the gas formation occurred only from the nanodroplets. In other words, the bubbles are not plasmonic nanobubles, which are transient vapor nanobubbles generated in surrounding liquid around laserFigure 9.Temperature change as a function of laser irradiation duration at 10 kHz overheated plasmonic nanoparticles pulse (at average power, 1.9 W). 28 . This is important because it suggests that the brightness on ultrasound imaging only comes from the nanodroplets phase-transitioning, which allows us to quantify drug release. Furthermore, bare nanodroplets without gold nanorod coating did not show phase-transition (Supporting Info, Figure S5). This is because direct laser energy absorption of the bare nanodroplets is not sufficient to cause vaporization 8. Thus, thermal energy converted from photon energy by the gold nanorods (plasmon resonance) vaporize the nanodroplets. Surface coverage by the gold nanorods on individual nanodroplets was calculated. In a bulk emulsion, it was found by UV-vis spectroscopy that about 58 % of gold nanorods added during synthesis attached to the nanodroplets. On an individual particle, the surface coverage ranged from 25% to 100% depending on the diameter, when it was 400 nm to 200 nm, respectively. With the average diameter of 300 nm, the results match well with the fraction of the positively-charged lipids on the nanodroplet surface coating, which is 0.5. This analysis, in addition to the zeta-potential results, implies that the binding mechanism of the negatively-charged gold nanorods on the nanodroplet surface is electrostatic interaction. The nanodroplets were very stable against aggregation and dissolution at 4°C. However, the nanodroplets at 37°C in DI water or in an isotonic solution experienced dramatically decreased stability. The nanodroplets at 37°C gradually showed phase-transition to bubbles, resulting in disappearance of the nanodroplets over time. The boiling point of the liquid core is reported as 29 ºC 29; however when the liquid is confined in a submicron-sized sphere structure, the boiling point increases theoretically due to surface pressure (Laplace pressure) 30. According to the Clausis-Clayperon equation, the boiling temperature of the nanodroplet with 300 nm in diameter is about 100 ºC. This explains the fact that the nanodroplets did not go through phase-transition for a week, staying the same size. Over time, when they aggregate and merge, it is possible that the boiling point of the large particles decreases to 37 ºC and the nanodroplets are no longer stable. Furthermore, we have noticed that mechanical forces such as pipetting, stirring and centrifugation also affected the stability. These phenomena are consistent with observations from other studies 31, demonstrating shear stresses due to pipetting cause the phase-transition. For clinical applications, the nanodroplets will not be affected by mechanical stirring or pipetting for the stability measurement. In addition, the stability in physiological conditions can be improved by using liquid for the core with a higher boiling point. When the nanodroplet undergo phase-transition, theoretically if the gas follows ideal gas law (pV=nRT), a volume expansion factor is 125, which corresponds to 5 time-increase in radius 32. We have observed that the radius increased approximately 5 times (analyzed by imaging software, ImageJ). Some bubbles 14 ACS Paragon Plus Environment

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looked bigger over time, still within a second after phase-transition, mostly because bubbles merged (Figure 4). Other studies have noticed the volume expansion factor is greater than 125 and explained the phenomena by intake of dissolved gases from the host fluid 32-33. Ultrasound imaging has been adopted mostly to qualitatively monitor the process of injection or to confirm the phase-transition from nanodroplets to bubbles 34-35. In this study, the area of brightness intensity from the ultrasound imaging results provides a promising analysis method to quantitatively measure the amount of drug release non-invasively. The in vitro data demonstrated that we can control a fraction of the nanodroplets to affect the cells. The fraction can be precisely controlled by the irradiation time. In addition, it implies that it is possible to release multiple times on-demand. For the nanodroplet concentration that we tested, 20 s of irradiation delivers approximately 30% of the drug. We note that the parameters can vary dependent upon the amount of nanodroplets applied, laser power, environment, etc. For example, with lower peak pulses, we have observed only several nanodroplets or none phase-transitioned. In addition, when we performed the phase-transition experiment in a water tank for ultrasound imaging, we have noticed the phase-transition efficiency was 5~10% compared to 30~35% through the air. Laser is an ideal tool to trigger drug release in the eyes because light passes through the lens well and can be precisely controlled 36-37.The ex-vivo results suggest the nanodroplet drug delivery system can be applied to posterior eye disease treatment by irradiating light through the lens and the drug release can be detected using ultrasound imaging. In addition, the area covered by drug increased after phase-transition due to the volume expansion of the nanodroplets, which is beneficial to treat vitreous seeds (tumors spread in the vitreous) with a single injection.

5. Conclusions Overall, we have developed nanodroplets that can be activated by light in a precise controlled manner to delivery drug to the cells. The activation can be monitored by ultrasound imaging quantitatively by analyzing ultrasound brightness intensity. Furthermore, our study has demonstrated the thermal effect by laser irradiation, which is critical to investigate before clinical applications to prevent any collateral damage to surrounding tissue. Lastly, the nanodroplets can be targeted for a potential use of targeted treatment.

Acknowledgement We thank Dr. Kevin Li in College of Pharmacy, University of Cincinnati (UC), for insightful discussion and resources. We also thank Dr. Todd Abruzzo in Cincinnati Children’s Hospital Medical Center and Dr. Zelia Correa in Department of Ophthalmology in College of Medicine, University of Cincinnati, for discussion about applications of the drug delivery system to retinoblastoma in clinical settings. We would like to acknowledge Chet Closson in Live Microscopy Core in College of Medicine, University of Cincinnati, for laser activation observation using a microscope, and Dr. Vesselin Shanov in Chemical Engineering, University of Cincinnati for allowing us to use his equipment, a Raman microscope and an IR camera. Lastly, we acknowledge CytoViva and Nikon companies for optical images. Funding Sources We would like to acknowledge start-up fund from University of Cincinnati. 15 ACS Paragon Plus Environment


Supporting Information Figure S1: A TEM image and a UV-vis spectrum of gold nanorods used in the study. Figure S2: XTT absorbance results of RPE (retinal pigment epithelial) cells at different concentrations of melphalan to determination of free melphalan concentration for in vitro drug delivery test. Figure S3: UV-Vis Absorbance (at 633 nm) of nanodroplet dispersions at different number densities. Figure S4: Captured image of ultrasound B-mode imaging of the nanodroplets in a test tube (Video S1) and quantitative analysis of the brightness. Video S1: Ultrasound B-mode imaging of the nanodroplets in a test tube. Table S1: Fluorescence data of doxorubicin release and % phase-transitioned. Text S1: Freeze-thaw effect on nanodroplet stability against rupture. Figure S5: Optical images of bare nanodroplets after laser irradiation.

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