For deep tissue imaging, fluorescent contrast agents with excitation and emission in the near-infrared (NIR) range are generally preferred as analytical and diagnostic tools compared with agents fluorescing in other regions, such as cyanine dyes (cy3 and cy5), green fluorescent proteins, and luciferin/luciferase reporter systems.1, 2 The advantages of NIR contrast agents result from greater depth of light penetration within this part of the spectrum. An incident NIR light beam penetrating a tissue surface is subject to less absorption and optical scattering than incident visible or infrared light .3 In addition, autofluorescence background from tissues is minimal in the NIR region. NIR imaging therefore presents the advantages of greater sensitivity over other imaging modalities, as well as greater imaging depths, allowing tissue visualization depths up to several centimeters.1 Optical imaging also eliminates exposure to ionizing radiation in contrast to conventional methods of cancer screening such as mammography or positron emission tomography. Therefore, NIR contrast agents, including organic dyes, quantum dots, and polymeric fluorophores, are attractive materials for biomedical imaging with applications in early cancer screening beyond the skin, particularly breast cancer.
Indocyanine green (ICG) is a Federal Drug Administration (FDA)-approved NIR dye that belongs to the polymethine class of NIR contrast agents.4 ICG is used to visualize blood flow and clearance, with applications in monitoring cardiac output, liver function, and neovascularization or macular degeneration.5, 6, 7, 8, 9, 10 The application of ICG imaging in early breast cancer detection has also been demonstrated based on the extravasation of ICG bound to blood proteins from leaky tumor vasculature.9, 11, 12, 13, 14, 15 Although ICG has an advantageous excitation and emission spectrum, with a of (em/ex) in water and (em/ex) when bound to serum proteins,16, 17 the use of ICG as a versatile contrast agent is limited by several properties. First, ICG is an unstable molecule susceptible to rapid aqueous, photo, and thermal instability.2, 5 For example, the half-life of ICG in aqueous solution is at room temperature in the dark and decreases significantly at physiological temperatures and with exposure to light.2 Degradation of ICG in aqueous solution is due to interactions with solvent radicals and ions leading to saturation of the carbon double-bonded chain and the formation of leucoforms (reduced, nonfluorescent forms of the dye molecule).2, 18 These leucoforms and other fragments are incapable of excitation and emission in the near-infrared portions of the spectrum. During light exposure, solvent radicals and ions along with radicals formed from photoexcited ICG accelerate the degradation process; when aqueous ICG is exposed to increased temperatures, the degradation process is accelerated in the same manner due to increased kinetics of radical formation.2, 18 Second, after administration, ICG exhibits nonspecific binding to blood proteins with a high affinity toward lipoproteins.5, 16 Finally, ICG is cleared rapidly from the body with a biphasic plasma clearance consisting of a fast initial half life of and a slow half life above one hour at low concentrations.2, 5, 19 Thus, formulations of ICG that provide increased chemical stability, protection from nonspecific protein binding, and enhanced circulation time are critical for expanding the repertoire of potential applications for this dye.
Polymeric nanoparticles (defined as in size) have been used as delivery vehicles for various therapeutic and imaging applications.20 Compared to free drug or imaging agents, nanoparticulate formulations present several advantages, including targeting capabilities and protection of therapeutic cargo against degradation or rapid clearance from the body. Polymeric nanoparticles passively target solid tumors due to the increased endocytotic activity, porous vasculature, and poor lymphatic drainage characteristic of these tissues, a phenomenon known as the enhanced permeation and retention (EPR) effect.21, 22, 23 Additionally, polymeric carriers can be functionalized to include active targeting ligands or to prolong circulation time.21 Recently, Saxena incorporated ICG into poly(lactic-co-glycolic acid) (PLGA) nanoparticles and demonstrated significantly enhanced photo, thermal, and aqueous stability of ICG compared to free ICG.24 In particular, encapsulation of ICG within the nanoparticle resulted in a degradation half life 4.3 times greater than that of free ICG when observed over a four day period at in the dark.24 In a different study, Yu developed silica/polymer capsules containing ICG for photothermal applications.25 These formulations also imparted photostability to encapsulated ICG. However, both of these ICG carriers would likely have limited utility for imaging of metastatic cancer due to particle sizes that are near or greater than the upper limit of fenestrations observed in tumor vasculature (mean diameter of approximately for the PLGA nanoparticles and for the silica/polymer capsules).24, 25, 26
The goal of our work is to develop a nanoparticulate delivery system for ICG that provides efficient dye loading, stabilized fluorescence over long time periods, potential chemistries for functionalization, and dimensions amenable to systemic delivery. The development of such a system will greatly improve the utility of ICG for early detection of breast cancer by increasing the stability of ICG and by providing a means to target ICG to tumor tissue. Here, we describe the synthesis, formulation, and characterization of ICG-containing polymeric micelles. Our results suggest that these materials protect ICG against degradation, without adversely impacting its spectral properties. Additionally, the micelle size and stability are ideally suited for passive targeting of ICG encapsulated micelles to the tumor milieu, and the polymer chemistry employed is readily adaptable for the incorporation of active targeting ligands.
Materials and Methods
All chemicals were purchased from Sigma-Aldrich (Milwaukee, Wisconsin) and used without further purification unless otherwise noted. The free radical initiator -azo-bis(isobutylnitrile) (AIBN) was recrystallized from methanol prior to use. Styrene was purified by distillation under reduced pressure. Maleic anhydride was recrystallized from chloroform. Solvents used were ACS grade and obtained from Sigma-Aldrich with the exception of ethyl ether (EMD, Gibbstown, New Jersey).
Poly(styrene-alt-maleic anhdyride)-block-poly(styrene) (PSMA-b-PSTY) copolymers were prepared by a thermally initiated, two-step reversible addition-fragmentation chain transfer (RAFT) polymerization using the chain transfer agent (CTA) S-benzyldithiobenzoate (BDTB). BDTB was prepared from S-(thiobenzoyl)thioglycolic acid and benzylmercaptan as previously described.27 In the first step, equimolar amounts of styrene and maleic anhydride were polymerized for at with AIBN as initiator and BDTB as chain-transfer agent in p-dioxane . The BDTB:AIBN ratio was 10:1, and the polymerization vessel was subject to three cycles of freeze-vacuum-thaw before polymerization. The PSMA macro-CTA (Mn , ) was isolated by precipitation in ethyl ether, filtered, and dried under vacuum at room temperature. In the second step, of macro-CTA were combined with styrene and AIBN (10:1 macro-CTA:initiator ratio) in dimethylformamide (DMF) in a round bottom flask. The solution was degassed by three cycles of freeze-vacuum-thaw, and polymerization was carried out for in a oil bath. The block copolymer PSMA-b-PSTY (Mn 21,600 kDa, ) was recovered by precipitation in ethyl ether, followed by filtration, and dried under vacuum at room temperature.
The molecular weights of the PSMA macro-CTA and PSMA-b-PSTY block copolymers were determined by size exclusion chromatography using Tosoh TSK-GEL and columns (Tosoh Bioscience, Montgomeryville, Pennsylvania) connected in series to a Viscotek GPCmax VE2001 and refractometer VE3580 (Viscotek, Houston, Texas). High pressure liquid chromatography (HPLC)-grade DMF containing LiBr was used as the mobile phase. The molecular weights of the synthesized copolymers were determined using a series of poly(methyl methacrylate) standards.
Following synthesis of the PSMA-b-PSTY block copolymer, the PSMA block was derivatized with butylamine as previously described,28 yielding a pH-sensitive hydrophilic PSMA derivative through aminolysis of 60% of the anhydride groups in the PSMA block. Briefly, a solution of PSMA-b-PSTY block copolymer was prepared in DMF. Subsequently, of a solution of butylamine in DMF was added drop-wise while stirring, and the solution was allowed to react for at room temperature. The polymer was recovered by precipitation in ethyl ether, and remaining anhydrides were hydrolyzed by dissolving the polymer in NaOH. After , the solution was extensively dialyzed against deionized water using SpectraPor Slide-a-Lyzer cassettes ( , Pierce, Rockford, Illinois) to remove sodium salts and neutralize the solution prior to lyophilization.
Modified PSMA-b-PSTY block copolymers were dissolved in DMF and NaOH . After of stirring, of nanopure (pH 9) was added by syringe pump at a rate of . Following injection, the solution was transferred to a dialysis cassette (MWCO 3500, Pierce) and dialyzed extensively against nanopure . Following dialysis, the solution was filtered through a filter.
Indocyanine Green Encapsulation
A solvent evaporation method was employed to encapsulate ICG into the micelles after complexing ICG with tetrabutylammonium iodide to form a hydrophobic ICG-tetrabutylamine salt. Briefly, a solution of ICG was prepared in chloroform with a six-fold molar excess of tetrabutylammonium iodide. The solution was sonicated for , then diluted in chloroform to obtain a final ICG concentration of . The chloroform ICG solution was then added to a stirring micelle solution , yielding a final ICG concentration of ICG. The chloroform was removed by evaporation, causing ICG to partition into the hydrophobic cores of the PSMA-b-PSTY micelles. Free ICG was then separated from the ICG-micelle solutions using Amicon regenerated cellulose centrifuge filters (MWCO , Millipore, Billerica, Massachusetts). Following ultrafiltration, the micelle-ICG solution was rinsed three times using nanopure . The carryover volume between rinses was less than . Purified micelles were resuspended in the original volume of nanopure . The final micelle concentration after filtering was determined by . Lyophilized samples were resuspended in deuterated DMF containing 1% trimethylsilane (TMS). Concentration was determined by comparing the styrenic peak area to a series of known standards using the TMS peak as an internal reference standard.
Indocyanine green loading efficiency
The efficiency of ICG loading was determined by lyophilizing samples of micelle-encapsulated ICG solutions. After drying, the micelle solutions were dissolved in dimethylformamide (DMF), causing complete dissolution of the micelle and release of the encapsulated ICG. The ICG concentration was determined by comparing ICG absorbance at to a standard curve of free ICG dissolved in DMF. The loading efficiency was then determined by the ratio of micelle-associated ICG to total ICG. The total ICG mass was defined as the amount of ICG initially added to the micelle solutions during encapsulation. All loading measurements were performed in triplicate.
Size, charge, and morphology of micelles
The size distribution of the micelles was determined before and after ICG loading by dynamic light scattering (DLS) in water and phosphate buffered saline (PBS , NaCl concentration of ) using a Brookhaven (Holtsville, New York) BI90Plus instrument equipped with a 535-channel correlator. A laser source was used as the incident beam, and measurements were performed at a angle. Calculations of particle diameter distributions were performed by the method of cumulants analysis. For unloaded and loaded micelles in PBS , the size was also examined after dilution in buffer to observe micelle salt stability. Zeta potential analysis was also performed on unloaded and loaded micelles in water using the Brookhaven BI90Plus. Atomic force microscopy (AFM) was used to determine micelle morphology and confirm the micelle size. A dilute solution of micelle-encapsulated ICG in water was applied on the surface of a cleaved mica disk (Ted Pella, Incorporated, Redding, California) and allowed to dry overnight. The mica surface was then rinsed with water and dried with nitrogen gas prior to AFM analysis. AFM scans of the micelle-coated mica surface were recorded and analyzed on a digital multimode AFM with Nanoscope IIIa controller (Veeco Instruments Incorporated, Woodbury, New York).
Critical micelle concentration
The critical micelle concentration (CMC) is the minimum concentration at which monomeric polymers dispersed in solution organize to form micelles. The CMC of the PSMA-b-PSTY system was investigated by observing the solvatochromic shift in fluorescence emission of 6-propionyl-2-(dimethylamino)naphthalene (PRODAN, Molecular Probes, Eugene, Oregon) as a function of block copolymer concentration. Following previously described protocols,29 a solution of PRODAN dye at in methanol was aliquoted into a series of glass vials. The vials were protected from light and dried under vacuum. Next, PSMA-b-PSTY solutions were added to the dry vials in triplicate at concentrations ranging from and allowed to stand overnight in the dark. The next day, the solutions were transferred to a 96-well polystyrene plate for fluorescence assays (Costar, Corning Incorporated, Lowell, Massachusetts) and read for PRODAN emission with excitation at ( bandwidth) and emission scan from ( bandwidth) on a Saphire 2 fluorescent microplate detector (Tecan, Austria). The CMC was determined by plotting the ratio of the peak hydrophobic emission intensity to the peak hydrophilic emission intensity against the polymer concentration on a logarithmetic scale. From this plot, the CMC was determined as the concentration at which the curve of the hydrophobic/hydrophilic ratio begins to increase with polymer concentration.
Formulation and Stability
Aqueous and thermal stability studies
The solution stability and thermal stability of ICG encapsulated within polymeric micelles were determined by assessing ICG emission over a three week period. For each reading, micelle-encapsulated ICG solutions were plated in triplicate into a 96-well plate. The emission intensity of the encapsulated ICG from (bandwidth ) was measured on the Tecan microplate detector using an excitation of (bandwidth ). All settings, such as the gain, integration time, number of reads, read position on plate, z-position of the scanner, and temperature, were maintained constant for the duration of the study. The fluorescence emission of ICG was recorded frequently within the first , approximately every for the next two weeks, and every within the third week. Controls consisted of free ICG in water prepared from the same stock as the encapsulated ICG, and unloaded micelles. For storage between readings, solutions were sealed in a glass vial and maintained in the dark at room temperature, or in a oven for the solution and thermal stability study, respectively. For analysis, normalized ICG emission was determined using the maximal emission intensity within a spectral bandwidth for each sample at a given time compared to the corresponding ICG emission at time zero. The average normalized emission intensity for each triplicate sample set was plotted as a function of time.
Kinetics of indocyanine green release from micelles
Release of ICG from freshly loaded polymeric micelles was examined for a period of . First, samples of ICG-encapsulated micelle solution were added to 25K MWCO dialysis tubing (SpectraPor, Spectrum® Labs, Incorporated, Rancho Dominguez, California), and the samples were dialyzed against of . During the study, the samples were stored in the dark at . At specified time points, fractions were collected from the solution within the dialysis tubing, and the remaining ICG concentration was determined by examining the absorbance of ICG at . A control consisted of free ICG. All measurements were performed in triplicate. The fraction of release was determined by comparing the ICG concentration found in each sample to that initially loaded.
In Vitro Study—Cytotoxicity and Value Determination
The (concentration of material that induces 50% cell death) of PSMA-b-PSTY micelles encapsulating ICG was determined using the MTS tetrazolium salt (3-(4,5-dimethylthiazol-2-yl)-5-(3-carboxymethoxyphenyl)-2-(4-sulfophenyl)-2H-tetrazolium salt) metabolic assay on NIH-3T3 cells. Cells were seeded at 12,000 cells/well in a 96-well tissue culture polystyrene plate (Costar Corning Incorporated Lowell Massachusetts). The cells were allowed to adhere to plates overnight prior to addition of micelle solutions. Stock micelle solutions at were sterilized by filtration through a syringe filter, diluted in cell culture media to final concentrations of , 0.6, 1.1, and , then added to cells. The cells were incubated in the presence of micelle solutions for in the incubator. The media was then replaced with fresh media, and the cells were returned to the incubator for . At the conclusion of the incubation, the cells were rinsed and treated with MTS reagent following manufacturer’s specifications (CellTiter 96® MTS Reagent Powder, Promega Corporation Madison, Wisconsin). Cell viability was evaluated by recording the absorbance of each well at using a microplate reader. Positive controls consisted of cells exposed to the same water/media concentration as each micelle treatment, and a negative control consisted of cells exposed to branched-polyethylenimine (b-PEI, MW 25K, Sigma, Saint Louis, Missouri).
Polymer Synthesis and Micelle Characterization
Block copolymers of poly(styrene-alt-maleic anhydride)-block-poly(styrene) (PSMA-b-PSTY) were successfully prepared with controlled architectures and low polydispersities using a two-step RAFT polymerization. Following polymerization, the maleic anhydride groups of the PSMA block were subject to ring-opening aminolysis using butylamine to obtain an amphiphilic copolymer as previously described (Fig. 1 ).28 60% of the maleic anhydride residues were opened by aminolysis with butylamine, followed by base-catalyzed hydrolysis to open the remainder of the maleic anhydride moieties. These modifications yield a block copolymer with a hydrophobic poly(styrene) component and hydrophilic poly(styrene-alt-maleic anhydride) component. The copolymers self-assemble in aqueous environments to form polymeric micelles with hydrodynamic diameters of as determined by dynamic light scattering. In addition, the overall charge of the micelles was determined to be near-neutral by zeta potential measurements (Table 1 ).
CMC, zeta potential, and particle sizing of unloaded and ICG-loaded micelles.
|Polymericmicelle state||CMC (mg/L)||Zeta potential(mV)||Solvent forsizing||Size (nm, freshlyprepared)||Size (nm, 50min after prepared)|
Indocyanine Green Encapsulation and Characterization
ICG was converted to the hydrophobic tetrabutylamonium salt and loaded into the hydrophobic micelle core by solvent evaporation. The efficiency of ICG loading was determined by computing the ratio of ICG concentration after loading to that of total ICG loaded, and was found to be . Next, the absorbance and fluorescence spectra of ICG were compared to the spectra of micelle-encapsulated ICG [Figs. 2a and 2b , respectively]. Slight red-shifts were observed in both the absorbance and emission spectra of encapsulated ICG (25 and , respectively).
The average diameter of ICG-loaded micelles was determined to be by dynamic light scattering. The overall charge of ICG-loaded micelles remained near-neutral (zeta potential of ). Particle sizing experiments revealed excellent salt stability of the empty and ICG-loaded micelles with minimal size change observed after of incubation in phosphate buffered saline containing NaCl. The particle characterization data are reported in Table 1. Micelle size and morphology were also determined by AFM after ICG loading. The AFM imaging of these particles showed an average diameter of (Fig. 3 ). Atomic force microscopy also revealed an overall spherical morphology, though it appeared the micelles flattened on the mica surface with the height of each individual particle smaller than its width along the surface. Based on the combined AFM and DLS data, it is reasonable to speculate that the micelles deform during the drying process before AFM imaging.
The critical micelle concentration (CMC) is an indication of micelle stability, a critical parameter for micelle-based therapeutics. Generally, the lower the CMC, the more stable the micelle. The CMC of the PSMA-b-PSTY micelles was determined both before and after loading ICG by measuring the solvatochromic shift of PRODAN fluorescence emission as a function of polymer concentration (Fig. 4 ). Above the CMC, PRODAN partitions from aqueous solution into the hydrophobic cores of polymeric micelles, undergoing a significant blueshift relative to its emission in water.29 Based on this method, the CMC of the PSMA-b-PSTY micelles alone was determined to be . The CMC for the ICG loaded micelles was also determined to be ; thus, ICG loading does not affect the CMC of PSMA-b-PSTY micelles (Table 1).
Solution Stability, Thermal Stability, and Release of Encapsulated Indocyanine Green
To determine whether encapsulation in polymeric micelles could protect ICG from degradation, the aqueous stability of ICG in micelle formulations was investigated. The fluorescence emission of micelle-encapsulated ICG and free ICG was examined over a period of three weeks with measurements repeated approximately every for the first two weeks and every within the third week. The emission profile of ICG over time provides a strong indicator of its overall stability in a given environment, and is a key parameter for imaging applications. In these studies, the fluorescence emission of free ICG decreased rapidly over time, while the fluorescence emission intensity of micelle-encapsulated ICG remained constant throughout the two week duration of the study (Fig. 5 ). The emission of free ICG decreased to 50% of its original value within , and was reduced by more than 90% within . These findings are consistent with the rapid instability of free ICG reported in the literature.2, 24, 25 In contrast, no change in ICG emission was observed when ICG was encapsulated in the micelle core.
Free ICG also degrades rapidly at elevated temperatures, another significant limitation for in vivo imaging applications. To evaluate the thermal stability of ICG-encapsulated polymeric micelles compared to free ICG, fluorescence emission of ICG at physiological temperatures was recorded from samples stored in a oven over a period of three weeks. Emission of ICG was recorded at the same intervals as in the solution stability study. At , the emission of free ICG decreased rapidly to 17% within and 0% within six days (Fig. 5). In contrast, ICG encapsulated in micelles was relatively protected from thermal degradation, maintaining 97% of original emission for . In addition, after three weeks of incubation at , encapsulated ICG still showed 41% of its initial emission. These results demonstrate the stabilization of micelle-encapsulated ICG at both room temperature and physiological temperature.
For in vivo imaging applications, the micelles must retain ICG until they have reached the appropriate targeted tissue, and must protect ICG from degradation and clearance until imaging is complete. The release of ICG from polymeric micelles was examined by measuring the retention of ICG within dialysis tubing as a function of time. Solutions of micelle-encapsulated ICG and free ICG were dialyzed against a MWCO 25K membrane in deionized water for at . At various time points, solutions were removed from the tubing, weighed, and concentrations of ICG remaining determined by absorbance readings at . By determining the concentration of ICG remaining in the dialysis tubing at a given time point relative to the initial absorbance, the mass loss of ICG was calculated as a percentage of the initial loading (Fig. 6 ). After , more than of the micelle-encapsulated ICG was retained in the tubing (i.e., 11% release), compared to remaining of free ICG (i.e., 73% release). This result indicates that the rate of ICG release from the micelles is sufficiently slow for in vivo imaging applications.
Cytotoxicity and Value Determination
A cytotoxicity study was performed to determine cell viability when exposed to increasing concentrations of micelle-encapsulated ICG. The , or concentration of polymer resulting in 50% cell death, was determined as an indicator of the toxicity of the polymeric micelles. Cell viability was determined by the MTS metabolic activity assay. No significant cytotoxicity compared to control treatments was observed at concentrations of 0.3, 0.6, and of PSMA-b-PSTY. Increasing toxicity with polymer concentration between (the highest concentration provided) was observed with cell viability decreasing from 88 to 33%, respectively, indicating an of PSMA-b-PSTY (Fig. 7 ).
Polymer Synthesis and Characterization
Previous studies have demonstrated the feasibility of one-pot synthetic procedures to prepare PSTY-b-PSMA block copolymers for the formation of micelles and cross-linked nanoparticles.30, 31 However, a two-step synthesis was employed here for ease of characterization of the molecular weight and polydispersity of each block in the copolymer. These block copolymers self-assemble to form polymeric micelles. Advantages of these specific micelles include the PSTY glassy core that contributes to micelle stability (low CMC), and the presence of functional groups in the corona that can be used for future derivitization with targeting ligands. The formulated micelles also have ideal sizes for tumor imaging applications. The desired size of nanoparticles for delivery from leaky tumor vasculature is less than .26 In addition, particles less than have decreased recognition by the reticuloendothelial system, resulting in longer circulation half lives.32 Prolonged circulation and reduced clearance will improve passive targeting to tumor tissue via the EPR effect. Unloaded micelles have an average diameter of , whereas ICG-loaded micelles have an average diameter of . The difference in size for the ICG-loaded micelles as opposed to the unloaded micelles is due to the process of ICG loading, not filtration, as was discovered in a separate study (data not shown). It is possible that the presence of chloroform during ICG loading allows reorganization of the micelle core. In physiologic salt concentrations, the sizes of both unloaded and loaded micelles were shown to be stable. The decrease in micelle hydrodynamic diameter after salt addition may be attributed in part to the salt shielding the anionic charges of PSMA polymers in the micelle corona, thereby decreasing electrostatic repulsion between the PSMA chains. In an in vivo setting, the micelles should be capable of maintaining their overall structure and thus retain cargo for circulation to the targeted site.
Indocyanine Green Encapsulation and Characterization
Tumor imaging using ICG is limited by the poor biodistribution and instability of the free dye. Recently, micelles have garnered great interest as carriers for therapeutic molecules because of their ability to improve the biodistribution and retention of small molecules, and their ability to protect fragile molecules during systemic circulation.32 By loading ICG into the cores of polymeric micelles, we hypothesized that the dye would be protected from rapid aqueous and thermal degradation without compromising the attractive spectral properties of the dye. To investigate this hypothesis, micelles derived from modified PSMA-b-PSTY block copolymers were loaded with ICG using a solvent evaporation method, which allowed ICG to partition into the hydrophobic micelle core. Using this method, ICG loading efficiencies of 87% were obtained from ICG loading solutions.
In aqueous solutions, free ICG is known to self-quench at concentrations above .12, 33 Therefore, self-quenching of ICG in the micelle core was a potential concern. To investigate the effects of ICG concentration on ICG fluorescence in the polymeric micelle, various concentrations of ICG were loaded into the micelles, and the fluorescence emission intensity was compared (data not shown). These studies showed micelle-ICG fluorescence increased as the ICG concentration in the loading solution was changed from , but decreased with further increases in ICG concentration. Based on these results, an initial ICG loading of into polymeric micelles was taken as the optimal concentration.
The ICG loading efficiency obtained in our studies compares favorably with previously published results. In other systems, drug loading efficiencies in polymeric micelles have been reported between 10 and 90%.34, 35, 36, 37, 38 In this study, the hydrophobic tetrabutylamine salt of ICG was loaded with 87% efficiency. A recent study by Saxena examined the stability of ICG when incorporated into PLGA nanoparticles, and calculated a loading efficiency.24 Also, Yu reported loading efficiencies between 90.0 and 97.1% when ICG was incorporated into much larger polymer nanoparticle-assembled capsules.25 In comparison to these previous studies, we can conclude that the ICG loading method we describe is highly efficient. Unlike previous investigations with ICG, we used tetrabutylammonium iodide to increase the hydrophobicity and micelle loading of ICG. The amount of tetrabutylammonium iodide used to form the ICG salt is 2300-fold below the (oral in rat is ), and should not pose complications for in vivo imaging applications.39
The stability of the polymeric micelles was determined by measuring the CMC before and after encapsulating ICG. Highly stable polymeric micelles (those with low CMCs) are required for in vivo applications, as micelles are subjected to significant dilution following intravenous injection.32 If injection reduces the block copolymer concentration below the CMC, polymeric micelles become thermodynamically unstable and begin to dissociate into component unimers. Dissociation is a kinetic process that occurs at different rates depending on the plasticity of the micelle core and other factors, but complete dissociation results in a loss of the therapeutic cargo. The CMC of the polymeric micelles described here was determined to be for both unloaded micelles and ICG loaded micelles. This value is lower40 or comparable41, 42 to that reported by other micelles derived from various amphiphilic block copolymers. The stability of the polymeric micelle system we describe is a result of the high glass transition temperature of the styrenic core, as has been well documented in previous work.32 For dosing in a mouse model, typical administration of ICG ranges from body weight.19, 43 This value corresponds to of micelles per kg or blood volume (assuming mouse blood volume is 5 to 6% body weight). Therefore, the target dose will be above the CMC of by almost three orders of magnitude.
Solution Stability, Thermal Stability, and Release of Encapsulated Indocyanine Green
A recent study examined the stability of ICG when incorporated into PLGA nanoparticles.24 ICG encapsulated in PLGA nanoparticles was stable after a four-day incubation in distilled water, showing 60% decrease in ICG fluorescence compared with free ICG (97.8% decrease).24 In the system presented here, micellar formulations of ICG resulted in sustained stabilization of ICG based on peak fluorescence for more than two weeks without any significant degradation. This reflects the polymeric micelles’ ability to stabilize ICG in an aqueous environment, thus allowing for easier formulation, longer shelf life, and a greatly enhanced diagnostic/therapeutic window. The micellar formulations of ICG also exhibited stability at ; fluorescence intensity decreased by only 59% after three weeks incubation in solution at physiological temperatures.
Over time, ICG is expected to release from the micelles as it diffuses out of the polystyrene cores. However, because of the glassiness of the poly(styrene) cores, we anticipated that the rate of ICG release would be very slow. The actual release rate of ICG from the micelle cores was determined to be 11% over . It should be noted that for the free ICG control, of ICG released as a burst from the dialysis tubing within the first six hours, followed by a period of slower release, resulting in a 73% total loss after . The plateau of release observed is likely due to fouling of the dialysis membrane by the released ICG. These results indicate the micelle-ICG formulation is sufficiently stable to allow for reasonable storage times and long imaging windows following in vivo administration.
In the present study, we show the encapsulation of ICG within polymeric micelles formed from poly(styrene-alt-maleic anhydride)-block-poly(styrene) (PSMA-b-PSTY) diblock copolymers. Characterization of the system shows efficient ICG loading, stabilized ICG fluorescence over varied conditions and long time periods, and minimal cytotoxicity. The polymeric micelle is capable of long-term retention of ICG, possesses a low CMC, and is readily adaptable for the incorporation of active targeting ligands. The PSMA-b-PSTY micelle system we discuss has the potential to greatly improve near-infrared imaging and detection of breast cancer by increasing the stability of ICG for formulation/administration, and by providing a means to target ICG to tumor tissue.
AFM studies were performed at the Nanotech User Facility, a member of the National Nanotechnology Infrastructure Network (NNIN), which is supported by the National Science Foundation and the Center for Nanotechnology at the University of Washington. We are grateful for the assistance of Xiangchun Yin, Adelaide Warsen, and Daniel MacDonald. This work was supported by an NSF Career Award (Li), the National Institutes of Health (R01 EB2991), University of Washington start-up funds (Pun), and a Ford Foundation Diversity Fellowship (Rodriguez).