Preparation and characterization of a polymeric (PLGA) nanoparticulate drug delivery system with simultaneous incorporation of chemotherapeutic and thermo-optical agents

Colloids and Surfaces B: Biointerfaces 75 (2010) 260–267

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Preparation and characterization of a polymeric (PLGA) nanoparticulate drug delivery system with simultaneous incorporation of chemotherapeutic and thermo-optical agents Romila Manchanda ∗ , Alicia Fernandez-Fernandez, Abhignyan Nagesetti, Anthony J. McGoron Department of Biomedical Engineering, Florida International University, 10555 West Flagler Street, Miami, FL 33174, USA

a r t i c l e

i n f o

Article history: Received 26 April 2009 Received in revised form 27 August 2009 Accepted 28 August 2009 Available online 4 September 2009 Keywords: PLGA Nanoparticles Indocyanine green Doxorubicin Fluorescence Solvent evaporation method

a b s t r a c t The objective of this study was to develop biodegradable poly(dl-lactide-co-glycolic acid) (PLGA) nanoparticles simultaneously loaded with indocyanine green (ICG) and doxorubicin (DOX). The modified oil in water single emulsion solvent evaporation method was used. To enhance the incorporation of both agents and control particle size, four independent processing parameters including amount of polymer, initial ICG content, initial DOX content, and concentration of poly-vinyl alcohol (PVA) were investigated. The ICG and DOX entrapment in nanoparticles as well as the nanoparticle size were determined. The nanoparticles produced by standardized formulation were in the range of 171 ± 2 nm (n = 3) with low polydispersity index (0.040 ± 0.014, n = 3). The entrapment efficiency was determined by spectrofluorometer measurements. The efficiency was 44.4 ± 1.6% for ICG and 74.3 ± 1.9% for DOX. Drug loading was 0.015 ± 0.001%, w/w, for ICG and 0.022 ± 0.001%, w/w, for DOX (n = 3). The release pattern was biphasic. ICG and DOX loaded-nanoparticle preparation was standardized based on the following parameters: PLGA concentration, PVA concentration and initial drug content. © 2009 Elsevier B.V. All rights reserved.

1. Introduction Indocyanine green (ICG) is a tricarbocyanine dye with infrared absorbing properties (peak absorption around 800 nm) and minimal absorption in the visible range. ICG exhibits an amphiphilic molecular structure with both hydrophilic and lipophilic properties, as shown in Fig. 1 [1,2]. ICG can dissolve in various solvents such as methanol, acetonitrile, DMSO, water, or PBS, but it has a tendency to aggregate and degrade rapidly in aqueous solution, creating challenges for its use in physiological media. ICG is FDAapproved and widely used for evaluation of cardiac output, liver and kidney function. It also has applications in photodynamic therapy, photothermal therapy and imaging [3–7]. ICG has been demonstrated to absorb near infrared (NIR) light at 808 nm (ideal for tissue penetration) and emit the energy as heat, which also makes it an ideal agent for localized hyperthermia with a rapid rate of temperature increase [8–10]. However, delivery of ICG for such uses in vivo is limited by its rapid plasma clearance, with a plasma half-life of about 3.2 min for free ICG [11]. This limitation can be overcome by incorporating ICG in carrier delivery systems for increased plasma stability.

∗ Corresponding author. Tel.: +1 786 399 2789; fax: +1 305 348 6954. E-mail address: rmanchan@fiu.edu (R. Manchanda). 0927-7765/$ – see front matter © 2009 Elsevier B.V. All rights reserved. doi:10.1016/j.colsurfb.2009.08.043

Doxorubicin (DOX) is an anthracycline drug (structure shown in Fig. 2) that is commonly used in the treatment of a large spectrum of solid tumors, such as breast and lung cancers [12]. Despite its efficacy at killing tumor cells, DOX has serious side effects, such as cumulative and irreversible cardiotoxicity, which limit the dosage and restrict treatment regimes [13]. Another area of concern with DOX treatment is the development of multidrug resistance in cancer cells, which leads to reduced efficacy of the treatments. These two issues have led researchers to investigate the development of advanced DOX delivery systems in order to provide improved therapies for the treatment of cancer [14]. One way to overcome the deficiencies of current treatment methods is to develop delivery systems that significantly improve the pharmacological characteristics of the drug in vivo, including specific targeting and increased delivery to sites of interest. In recent decades, much research has been focused on the application of biocompatible and biodegradable polymers to drug delivery systems [15–23]. Among those polymers, a significant one has been poly(dl-lactide-co-glycolide) (PLGA) from the ester family. This polymer has been widely used in the biomedical industry as a major component due to its biocompatibility and biodegradability [18,21,22]. PLGA polymers have been approved by the FDA and have been used in humans for many years as a resorbable suture material [24]. PLGA nanoparticles have also emerged as drug carriers in novel drug delivery systems because it is relatively easy to entrap hydrophobic drugs into PLGA nanoparticles, due to the hydropho-

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Fig. 1. Structure of ICG. ICG is an amphiphilic molecule. Two polycyclic parts interconnected by a carbon chain impart lipophilicity to the ICG molecule, whereas the sulfate groups are responsible for the hydrophobic nature.

bic nature of PLGA molecules. The incorporation of DOX into PLGA nanoparticles is a possible approach to reduce chemotherapy cytotoxicity to healthy tissues, which may decrease the undesirable side effects such as impaired cardiac function. There are numerous papers available in the literature reporting single-agent incorporation into PLGA nanoparticles [25–27], including the entrapment of ICG [28], as well as the entrapment of DOX [29]. However, PLGA nanoparticles loaded with dual agents were seldom reported [30,31]. Another approach to improve the efficacy and selectivity of cancer treatments is the application of hyperthermia in combination with traditional cancer therapeutics, such as radiation therapy and chemotherapy [32,33]. Hyperthermia makes some cancer cells more sensitive to radiation and it can also enhance the effect of certain anticancer drugs [32], thus allowing the use of decreased chemotherapy dosages. Hahn et al. reported a combined effect of hyperthermia with adriamycin chemotherapy in vivo and in vitro [34]. Numerous in vitro studies have shown the added cytotoxic effects of hyperthermia treatment in combination with chemotherapy [35,36]. Clinical studies have also demonstrated the efficacy of hyperthermia combined with radiotherapy and/or chemotherapy [33,37]. Liposomal doxorubicin combined with hyperthermia has been found to reduce toxicity and improve efficacy [38]. Our previous research also showed that simultaneous use of ICG and DOX in combination with localized hyperthermia can produce the same effect as that achieved by greater doses of chemotherapy alone [39]. Given the need for improved cancer therapeutics, the development of nanoparticulate delivery systems as drug carriers, and the proven combined effect of hyperthermia and chemotherapy, the objective of our present study was to develop a polymeric nanoparticulate delivery system containing a chemotherapy agent and an optical agent. The role of the optical agent is twofold. First, it serves as an imaging tracer to monitor drug delivery. Second, by selecting an optical agent that generates a significant enough amount of heat when exposed to a given wavelength of light, the molecule can be used as a hyperthermia source once the drug system reaches

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its target. DOX was selected as the chemotherapeutic agent, ICG as the optical tracer which can generate heat by absorbing NIR light, and PLGA nanoparticles as the carrier. As discussed previously, dual entrapment into PLGA nanoparticles has been rarely explored. In our study, the use of ICG and DOX constitutes a challenge, because these molecules are ampiphillic and hydrophobic, respectively. Therefore, dual entrapment into PLGA nanoparticles to obtain acceptable concentration levels of both agents requires the optimization of incorporation methods. Different methods of preparation of nanoparticles from preformed polymers have been reported in the literature, including solvent diffusion, emulsion evaporation, salting out evaporation, solvent displacement, and others [40]. Each method has its own advantages and limitations, so the appropriate choice of method is essential [41]. The method of solvent diffusion is limited by the need to use low polymer concentrations in order to obtain nanoparticles within a size range of 200 nm. Solvent evaporation is limited by its lengthy and expensive procedure. The salting out method is useful due to its greater sensitivity to high polymer concentrations, but it also has a large number of purification steps involved in the synthesis method. The emulsion evaporation method has the advantage that it can be used to entrap both hydrophobic and hydrophilic drugs by formulation of w/o and w/o/w emulsions, respectively. The ultimate choice of method for nanoparticle formulation can be based on chemical characteristics of the active component (the drug of choice) and its interactions with the organic solvents, polymer, surfactant, and the final use of the nanoparticles. For example, Saxena et al. [28] used methanol for dissolution of ICG in order to attain maximum retention of its fluorescence. Continuous efforts have been taken towards improvement of nanoparticle size, polydispersity index and entrapment efficiency and drug loading of both hydrophobic and hydrophilic drugs. These efforts have further motivated the research and discovery of new methods of nanoparticle formulation based on slight modifications of standard methods. In the present investigation, we have explored the potential of PLGA nanoparticles to simultaneously entrap ICG and DOX in order to create ICG–DOX nanoparticles (ICG–DOX–NPs). The PLGA nanoparticles were prepared by the oil in water single emulsion solvent evaporation methodology, and characterized for surface morphology and size employing scanning electron microscopy (SEM) and dynamic laser scattering (DLS). The influence of various processing parameters such as PLGA concentration, PVA concentration and drug entrapment on nanoparticle size was also investigated. Subsequently, the in vitro release kinetics of ICG and DOX from the nanoparticles was monitored at physiological buffer conditions employing phosphate buffer saline (pH 7.4). 2. Materials and methods Poly(dl-lactide-co-glycolide) (PLGA) of molecular weight 45–75 kDa with lactide–glycolide ratios 50:50 and poly-vinyl alcohol (PVA) with 88% hydrolyzation degree were purchased from Sigma–Aldrich (St. Louis, MO). Indocyanine green (ICG) and doxorubicin hydrochloride (DOX) were purchased from Fisher Scientific (Pittsburgh, PA) and Sigma–Aldrich, respectively. All other chemicals and solvents were of reagent grade (purchased from Sigma–Aldrich). 2.1. Preparation of nanoparticles

Fig. 2. Structure of DOX. Doxorubicin has two phenol groups and a daunosamine moiety with an ionizable NH2 group.

PLGA nanoparticles loaded with indocyanine green and doxorubicin were prepared by using a modified version of the oil in water single emulsion solvent evaporation process [30]. Briefly, PLGA polymer and the two drugs were dissolved in a methanol–dichloromethane mixture. This organic phase was

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emulsified with an aqueous phase of PVA solution by sonication using a probe sonicator (Cole Parmer) at an output of 50 W for 30 s in an ice bath. Nanoparticles were immediately formed. The organic mixture was then rapidly removed by evaporation under reduced pressure by using a Rotavapor R-205 (Buchi, Switzerland). The resulting PLGA nanoparticle suspension was then ultra-centrifuged at 14,000 rpm for 30 min. After centrifugation, the nanoparticle precipitate was washed by using the same volume of distilled water as the supernatant, and again centrifuged at 14,000 rpm for 15 min. The washing process was repeated three times in order to remove the adsorbed drugs. The washed nanoparticles were then freeze-dried using the Freezone system (Labconco, Free Zone Plus 6) for 24 h. The effect of various processing parameters and polymer characteristics on hydrodynamic particle size and percentage drug entrapment was studied. The processing parameters included polymer concentration in the organic phase, PVA concentration in the aqueous phase, initial ICG content, and initial DOX content. Unless otherwise mentioned, all the experiments were conducted by varying one of the parameters and keeping all the other process parameters constant, at a set of standard conditions as follows: 30 mg/ml of 45–75 kDa PLGA (L:G = 50:50), 10 ␮M of ICG and 13.3 ␮M of DOX prepared in a 1.5 ml mixture of dichloromethane:methanol (2:1, v/v) as the organic phase, and 4.5 ml of 2% PVA solution as the aqueous phase. The aqueous-toorganic phase volume ratio was 3:1. 2.2. Characterization of nanoparticles 2.2.1. Size determination and zeta potential measurement (i) Dynamic light scattering (DLS): the hydrodynamic diameter of the nanoparticles was determined by dynamic light

Nanoparticle yield (%) =

sonicator for 30 s to create a homogeneous suspension. The zeta potential was measured in triplicate using a Zetasizer, Nano ZS (Malvern instruments, UK).

2.2.2. ICG and DOX entrapment efficiency The freeze-dried nanoparticles containing ICG and DOX were weighed and dissolved in dimethylsulfoxide (DMSO). The fluorescence spectrum of the samples was measured using a Fluorolog-3 spectrofluorometer (Jobin Yvon Horiba) in steady-state mode. After a 30-min warming period, the instrument was calibrated following manufacturer specifications for both the xenon lamp source and the expected Raman spectrum of water. ICG and DOX calibration curves were prepared prior to measurements to verify spectral characteristics, linearity range, and degree of overlap between the spectra of the two drugs. An absence of overlap ensures specificity of the fluorescence readings, and eliminates the need for mathematical correction of the measurements. These calibration curves were obtained using the same instrument and operating conditions. Samples were placed in polystyrene 4.5-ml cuvettes and measurements were performed immediately after preparation. DMSO was used as a blank for background correction. The excitation wavelengths were 785 nm for ICG and 496 nm for DOX, and spectral emission readings were recorded in 1-nm intervals up to 870 nm. All measurements and sample handling were done in reduced lighting conditions, and the instrument operating conditions were kept constant. The concentrations of ICG and DOX were determined from the corresponding blank-subtracted reading at the peak emission wavelength, by using the previously obtained calibration curves of ICG and DOX in DMSO. The nanoparticle yield and entrapment efficiency were calculated as follows:

Mass of nanoparticles recovered × 100 (Mass of PLGA + Mass of ICG + Mass of DOX) in formulation

Drug (ICG or DOX) entrapment efficiency (%) =

scattering (DLS) measurements. Nanoparticles were suspended in 1 ml of double-distilled water and sonicated prior to measurements. Nanoparticle size was determined using a Zetasizer, Nano ZS (Malvern instruments, UK) employing a nominal 5-mW He–Ne laser operating at 633 nm wavelength. The scattered light was detected at a 135◦ angle. The refractive index (1.33) and the viscosity (0.89) of ultrapure water at 25 ◦ C were used for size measurements. The measured nanoparticle sizes were presented as the average value of 20 runs, and experiments were done in triplicate. The polydispersity index of the particle size was measured using the same instrument on a scale from 0 to 1. (ii) SEM: scanning electron microscopy (SEM) was used to verify uniformity of particle shape and size. Freeze-dried nanoparticles were resuspended in distilled water and were later dropped onto a silicon grid and dried under room temperature. The nanoparticle suspension was vacuum-coated with gold for 3 min. The surface morphology of the samples was observed under a scanning electron microscope (JEOL–JEM) operated at 15-keV pulse at different resolutions [42]. (iii) Zeta potential measurements: in order to measure the zeta potential of the nanoparticles, a dilute suspension of nanoparticles was prepared by resuspending 0.1 mg of freeze-dried nanoparticles in 1 ml of distilled water at pH 7.4. After resuspending in water, the particles were sonicated in a bath

Mass of drug (ICG or DOX) in nanoparticles × 100 Mass of drug (ICG or DOX) used in formulation

(1)

(2)

2.2.3. In vitro release kinetics Freeze-dried nanoparticles (20 mg) were resuspended in 40 ml of phosphate-buffered saline (PBS), pH 7.4 and were bath sonicated. For each of the two drugs, the total drug amount inside the nanoparticles was kept below 10% of the drug solubility limit in PBS in order to ensure sink conditions. The solubility limit of ICG in PBS is approximately 35 mg/ml, and for DOX the solubility limit is 50 mg/ml. The suspension was divided into aliquots in microcentrifuge tubes and the tubes were placed in an incubator oven (VWR Scientific, USA) at 37 ◦ C and stirring speed of 80 rpm. The centrifuge tubes were collected at regular time intervals and were centrifuged at 14,000 rpm for 20 min. The nanoparticles were then washed with the same amount of fresh buffer and again centrifuged at 14,000 rpm for 6 min followed by lyophilization for 36 h. Remaining drug (ICG and DOX) in the nanoparticles after release was measured by dissolving the nanoparticles in DMSO so that the drug contents are released into solution. Release kinetic studies were done up to 48 h from the original time point (t = 0 h). The cumulative percentage of ICG and DOX released from the nanoparticles at a certain time interval was calculated by the following equation: Cumulative % ICG or DOX released = 100 − (% ICG or DOX remaining)

(3)

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2.3. Statistical analysis The statistical analysis of the data was done using independent Student’s t-test between the groups on Statistical Product and Service Solutions (SPSS V 13.0). The significance level used was ˛ = 0.05. 3. Results and discussion 3.1. Effect of different parameters on formulation characteristics

PLGA nanoparticles with 10 ␮M ICG and 13.3 ␮M DOX as initial drug contents were also prepared.

Formation of an emulsion is the most important step in preparation of nanoparticles because the size of emulsion droplets is directly related to the final nanoparticle size. Emulsification involves mixing of an organic phase consisting of polymer with an aqueous phase containing a surfactant or stabilizer. This emulsion is broken down into droplets by applying external energy (through sonication), and these nano-droplets lead to nanoparticle formation upon evaporation of the organic solvent. The total volume and total viscosity are related to the effectiveness of mixing, whereas the stirring speed relates to the amount of energy put into the system. These are the crucial parameters to determine the final particle size. Table 1 displays the formulation parameters used in the standardization.

a

Formulation 5

N/A N/A ICG: 35 ␮MDOX:37 ␮M

Formulation 4

100 mg 5% ICG: 30 ␮MDOX: 32 ␮M

Formulation 3

80 mg 4% ICG: 25 ␮MDOX: 27 ␮M

Formulation 2

40 mg 2% ICG: 20 ␮MDOX: 22 ␮M

Formulation 1

2% PVA, 10.0 ␮M ICG, 13.3 ␮M DOX 45 mg PLGA, 10.0 ␮M ICG, 13.3 ␮M DOX 45 mg PLGA, 2% PVAa

20 mg 1% ICG: 15 ␮MDOX: 17 ␮M

Variable parameters

Amount of PLGA (mg) PVA Concentration (%) Initial drug (ICG and DOX content)

Constant parameters

Table 1 Formulation parameters for the standardization of ICG + DOX loaded PLGA nanoparticles.

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3.1.1. Effect of PLGA concentration on particle size and zeta potential Fig. 3 shows the effect of polymer concentration in the organic phase on nanoparticle diameter and on zeta potential. The PLGA concentration in the organic phase (Methanol:EDC, 1:2, v/v) was varied from 13.3 to 66.7 mg/ml, whereas all other processing parameters such as PVA concentration, sonication time, and drug concentration were kept constant. An increase in PLGA concentration causes a significant (p < 0.05) increase in nanoparticle diameter of the ICG–DOX–NPs. The mean size of the nanoparticles (n = 3) was 137 ± 2 nm for 13.3 mg/ml, and 164 ± 2 nm for 66.7 mg/ml. This increase in nanoparticle size with an increase in polymer concentration is in accordance with the results reported by other researchers [30,31,43,44]. Increasing the PLGA concentration in the organic phase increases the viscous resistance of the emulsion mixture, thereby absorbing the agitation energy which in turn leads to reduction in shear stress resulting in droplets with larger size. An increase in PLGA concentration causes a significant (p < 0.05) decrease in zeta potential value of the ICG–DOX–NPs. The mean zeta potential (n = 3) was −9.9 ± 0.4 mV for 13.3 mg/ml, and −12.3 ± 0.1 mV for 66.7 mg/ml. The decrease in zeta potential value

Fig. 3. This figure shows the change in particle size and zeta potential due to changes in PLGA concentration. PLGA nanoparticle size increased and zeta potential decreased with increasing PLGA concentration in the organic phase.

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with an increase in PLGA concentration is consistent with the results reported by Xie et al. in the preparation of solid lipid nanoparticles [45]. 3.1.2. Effect of PLGA concentration on entrapment efficiency Fig. 4 shows the effect of PLGA concentration on the entrapment of the two drugs. The entrapment of both drugs (ICG and DOX) increases with an increase in PLGA concentration. This increase in entrapment is due to an increase of viscosity in the organic phase, which results in an increase in the diffusion resistance of drug molecules from organic phase to the aqueous phase. Furthermore, the increase of particle size with the increase in polymer concentration can increase the length of diffusional pathways of drugs from the organic phase to the aqueous phase, thereby reducing the drug loss through diffusion and increasing the entrapment. Other groups [27,31] have also reported an increase in their drug entrapment by increasing PLGA concentration due to similar reasons. Additionally, increasing PLGA concentration creates a larger number of polymer matrices for the entrapment of drug molecules. This is probably the reason for such a drastic increase (p < 0.05) in the entrapment of DOX molecules due to their hydrophobic nature. Entrapment efficiency for ICG showed a gradual increase (p > 0.05) with increase in PLGA concentration. The entrapment of ICG first increases and then reaches a plateau. These results are in accordance with the results obtained by Saxena et al. and Gomes et al., who reported that PLGA nanoparticles have a fixed capacity to entrap highly amphiphillic ICG molecules [28,46]. 3.1.3. Effect of PVA concentration on particle size and zeta potential Stabilizer addition also plays an important role in the formulation of nanoparticles and in their resulting size. The role of the surfactant molecule is to stabilize the emulsion droplets by preventing them from coalescing. The surfactant molecules have to cover the interface between the organic solvent and the aqueous phase for all the droplets, in order to obtain effective stabilization. Therefore, the properties and concentration of stabilizer will change the total surface area of the nanoparticles, and will have an effect on the final particle size and entrapment efficiency. For our studies, we used PVA as a stabilizer. The concentration of this stabilizer was varied to identify its effect on particle size and entrapment efficiency, keeping the other parameters constant as in the standard procedure. Fig. 5 shows the effect of PVA concentration in the aqueous phase on nanoparticle size and on zeta potential. With increasing PVA concentration (1–5 wt.%, n = 3), the particle size decreases significantly (p < 0.05) from 159 nm (1% PVA) to 113 nm

Fig. 5. This figure shows the effect of PVA concentration on PLGA nanoparticle size and zeta potential. Note the decrease in PLGA nanoparticle size and the increase in zeta potential with increasing PVA concentration.

(5% PVA). Some researchers have reported a decrease in particle size with increased stabilizer concentration [25,31,44]. They suggest that, at higher concentrations, more PVA can be oriented at the interfacial area (organic solvent/water) so that the interfacial tension is reduced [47]. This results in an increase in the net shear stress at constant energy density during emulsification, leading to the formation of smaller droplets and smaller nanoparticles. However, contradictory results have been reported by other groups [48,49]. They report that, by increasing the PVA concentration, the viscosity of the external aqueous phase increases, which results in a decrease in net shear stress and a corresponding increase in particle size. For our system of PLGA nanoparticles, reduction in interfacial tension may be the dominating factor so that we observe a decrease in particle size with increasing PVA concentration. With increasing PVA concentration (1–5 wt.%, n = 3), the zeta potential increases significantly (p < 0.05) from −23.6 ± 1.8 mV for 1% PVA to −10.4 ± 1.7 mV for 5% (w/v) PVA. Stolnik et al. reported a negative zeta potential of −45 mV for PLGA nanoparticles without PVA in neutral buffer [50]. This high negative surface zeta potential of the nanoparticle can be attributed to the uncapped carboxyl groups of the PLGA polymer [51]. Previous studies [51–53] have shown that nanoparticles prepared using PVA still have negative zeta potential, but it shifts towards the positive direction. This observation is consistent with our results. The reason for this shift may be that a high PVA concentration increases the number of coating layers on the polymer surface and shields the surface charge of the particles. 3.1.4. Effect of PVA concentration on entrapment efficiency Fig. 6 shows the entrapment efficiencies of ICG and DOX. The entrapment efficiency of the two drugs decreases significantly with an increase in PVA concentration (p < 0.05). This decrease in entrapment of drugs with an increase in PVA concentration is in accordance with the results reported in the literature [30]. The change in entrapment efficiency with PVA concentration is mainly a result of changing particle size. Also, by increasing PVA concentration, more molecules of the two drugs may partition out rapidly into the aqueous phase during the emulsification procedure, and fewer drug molecules remain in the emulsion droplets to interact with PLGA molecules, hence decreasing the entrapment efficiencies [54].

Fig. 4. This figure shows the change in entrapment efficiency of ICG and DOX with respect to change in PLGA concentration. The entrapment efficiency of ICG and DOX increased significantly with increasing PLGA concentration in the organic phase.

3.1.5. Effect of initial drug content on particle size and drug entrapment efficiency Fig. 7 shows that the particle size is independent of the initial ICG and DOX concentration in the organic phase (p > 0.05), which is consistent with existing literature [27,30,31].

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Fig. 6. This figure shows the effect of PVA concentration on the entrapment efficiency of ICG and DOX in PLGA nanoparticles. The entrapment efficiency of both drugs decreased significantly with increasing PVA concentration.

Fig. 8a and b show the effect of initial drug concentration of ICG and DOX on the entrapped drug concentration. As the initial drug concentration of both drugs increases, the entrapped drug concentration first increases and then reaches a plateau. The maximum entrapped drug concentration is approximately 4.3 ␮M for ICG (at initial drug concentration of 30 ␮M), and approximately 5.9 ␮M for DOX (at initial drug concentration of 22 ␮M). The drug content in the nanoparticles is affected by drug–polymer interactions and by the drug miscibility in the polymer. The importance of drug miscibility in the polymer has been discussed by Panyam et al. [55] for a hydrophobic drug–polymer system of dexamethasone or flutamide-loaded PLGA/PLA nanoparticles. This group reported that higher drug–polymer miscibility resulted in higher drug incorporation. In our experiments, with an increase in initial drug content during formulation, the drug concentration in the organic phase increases. This means that more drug molecules can interact with PLGA molecules, resulting in an increase of the entrapment amount and entrapped concentration of ICG and DOX. However, the increase in entrapment amounts is not in proportion to the increase in initial drug content during formulation, so the entrapment efficiencies decrease. 3.2. Standardization of ICG–DOX–PLGA nanoparticles Previous research has shown that concentrations of 5 ␮M ICG and 10 ␮M DOX are effective in producing hyperthermia in combination with chemotherapy [39]. Considering these results,

Fig. 8. (a) This figure shows the concentration of ICG entrapped in the PLGA nanoparticles as a function of initial molar concentration of ICG. For some data points the standard deviation is small (less than 1%), but all data points include error bars. Entrapped ICG concentration in the nanoparticles increases with increasing initial molar concentration in the formulation. (b) This figure shows the concentration of DOX entrapped in the PLGA nanoparticles as a function of initial molar concentration of DOX. For some data points the standard deviation is small (less than 1%), but all data points include error bars. Entrapped DOX concentration in the nanoparticles increases with increasing initial molar concentration in the formulation.

along with the fact that high drug entrapment efficiency, high entrapment concentration, and small particle size are desirable features of our drug delivery system, we standardized the preparation of ICG–DOX–PLGA nanoparticles. Standardized ICG–DOX–PLGA nanoparticles were prepared by taking 80 mg of PLGA, 20 ␮M ICG and 22 ␮M DOX, and dissolving in 1.5 ml of dichloromethane:methanol (2:1, v/v) mixture. This organic phase was emulsified with 4.5 ml of PVA solution (1%, w/v) by probe sonication at 50 W for 1 min in an ice bath. The organic solvent was then rapidly evaporated under reduced pressure. 3.3. Characterization of standardized nanoparticles

Fig. 7. Effect of initial drug concentration on particle size. Initial drug concentration did not have any significant effect on PLGA particle size. The x-axis labels “−/−” indicate ICG/DOX concentrations in ␮M.

Particle size and surface morphology of ICG–DOX–PLGA nanoparticles were determined by employing DLS and SEM (n = 3). Nanoparticle size was 171 ± 2 nm diameter as determined by DLS (Fig. 9). The polydispersity index of 0.040 indicated a narrow size distribution. In SEM pictures, the particles show diameters of about 90–100 nm. Fig. 10 shows a sample SEM image at 50,000× magnification. DLS measures the hydrodynamic radii by dispersing particles in aqueous phase or solvents, whereas SEM measures the size of dried samples loaded with a Pt/Pd mixture and then vacuum-dried with gold. We believe that the hydration and swelling of the particles in aqueous buffer may be the possible reason for observing

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from the organic phase into the aqueous phase would be larger than for haloperidol. 3.4. Release kinetics

Fig. 9. DLS after optimization of nanoparticle formulation. This is a representative overlay of dynamic light scattering of ICG–DOX–PLGA nanoplexes after standardization. The average size in this case was 171 nm.

Fig. 10. SEM image of ICG–DOX–PLGA nanoparticles after standardization of the formulation. The average particle size of the nanoparticles was approximately 100 nm. The image magnification is 50,000×.

larger sizes by DLS measurements as compared to SEM. Nanoparticle yield was 80.0 ± 2.0%. The entrapment efficiency was 44.4 ± 1.6% for ICG and 74.3 ± 1.9% for DOX. Drug entrapment efficiency for DOX was relatively lower than expected when compared with previous entrapment reports for lipophilic drugs such as haloperidol [27]. We believe that the difference can be explained by the degree of hydrophobicity of the two drugs, with haloperidol being much more hydrophobic than DOX. The solubility of DOX in water is 50 mg/ml, whereas haloperidol solubility in water is 0.014 mg/ml. Therefore, we would expect that the amount of DOX partitioned

Fig. 11 shows the biphasic pattern of drug release (ICG and DOX) from the PLGA nanoparticles. Drug release occurs in two phases, an initial “burst” release and a much slower diffusion-initiated release due to the concentration gradient. ICG shows a 56% release after 8 h, and a much slower release thereafter. Cumulative ICG release at 48 h is 72%. DOX Exhibits 48% release after 8 h, and a cumulative release of 51% at 48 h. As can be seen in Fig. 11, ICG release is much faster than DOX release. This agrees with the results from Kalaria et al. [56], who showed a slow DOX release from PLGA nanoparticles with 80% release at 24 days. Since a spectrofluorometer was used to measure drug concentration rather than an HPLC, we followed the drug release profiles for only 48 h. The results, however, provided important observations regarding the difference in release behavior between ICG and DOX. The hydrophilic ICG molecule shows a faster release profile than does the DOX, which is hydrophobic in nature. The incomplete cumulative release can be explained by a combination of the short time duration of the study as well as the increased initial drug content. Other groups have reported that increases in initial drug content result in increased polymer–drug interactions, which make the nanoparticle matrix more rigid, and cause incomplete and slower release of drugs from the nanoparticle [57,58]. The results of our release studies are consistent with those of other authors who have explored the in vitro release kinetics of different drugs entrapped in PLGA nanoparticles. Musumeci et al. reported that release of hydrophobic docetaxel from PLGA nanoparticles was between 40 and 68% within a 24-h sampling period, depending on drug formulation. Release of the drug after 10 days was between 70 and 95% [59]. Therefore we may expect similar release kinetics from hydrophobic drugs entrapped in PLGA nanoparticles. We observed 51% release of DOX at 48 h. For ICG-PLGA nanoparticles, Saxena et al. reported that 80% of ICG was released after 16 h and the release plateaus after this time point [28]. We observed 72% release of ICG at 48 h. In conclusion, we observed incomplete release of both DOX and ICG from the PLGA nanoparticles after 48 h, and this is consistent with existing literature on the individual behavior of these two drugs when entrapped in nanoparticle systems. In our experiments, we only studied “burst” and “diffusion” related releases. We have not accounted for other release mechanisms such as release due to osmolarity or swelling of nanoparticles

Fig. 11. This figure shows the release kinetics of ICG and DOX from the PLGA nanoparticles over a 48-h period. Note the biphasic release profile for both drugs, with an initial burst release followed by a slower diffusion-related release pattern. ICG release was faster and cumulatively higher than DOX release.

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[60], but the DLS measurements of PLGA nanoparticles after 48 h in suspension show that the nanoparticles have the same size distribution as in the initial measurements, indicating that particles are not likely to have swollen in the 48-h period and there is no evidence of PLGA bulk erosion. Additionally, literature results show that PLGA nanoparticles are stable in a suspension of PBS at 37 ◦ C and pH 7.4 for up to a period of 10 weeks with no significant change in size [23]. 3.5. Expected advantages of the dual agent approach Our approach has the advantage of combining localized hyperthermia, chemotherapy, and optical imaging into a single drug delivery vehicle. The ICG–DOX nanoparticles can be conjugated with functional groups and decorated with antibodies for targeted cell uptake, which will result in preferential accumulation at tumor sites and reduced systemic toxicity. The incorporation of both agents into one vehicle allows for simultaneous uptake of the two drugs into tumor tissue, which facilitates the combined hyperthermia–chemotherapy effect. Simultaneous delivery also has the potential to simplify drug administration protocols by eliminating the need to administer two different preparations and coordinate their uptake profiles at the target site. The fact that ICG can also serve as an imaging agent is an added advantage, because it will allow for enhanced localization of the tumor and non-invasive visualization. Potentially, the imaging component could allow for monitoring tumor size changes and could be used as an additional tool in treatment planning. 4. Conclusion In this paper, we used a modified method of oil in water single emulsion to produce PLGA nanoparticles loaded with dual ICG–DOX agents. Entrapment efficiency of drug was dependent on different formulation factors such as polymer concentration, PVA concentration and drug concentration. To increase drug loading in the nanoparticles, the amount of ICG and DOX in the formulation was varied from 10 to 35 ␮M for ICG and 13.3 to 37 ␮M for DOX, and the amount of polymer was varied from 13.3 to 66.6 mg/ml. Additionally, by changing the surfactant concentration from 1 to 5%, w/v, the drug uptake was standardized. Drug loading was 0.015 ± 0.001%, w/w, for ICG and 0.022 ± 0.001%, w/w, for DOX (n = 3). Both drugs showed incomplete release from the nanoparticles over a period of 48 h. We demonstrated that the modified oil in water single emulsion solvent evaporation method can be used to successfully and simultaneously entrap ICG and DOX in PLGA nanoparticles. The significance of this study is the synthesis of multi-functional polymer nanoparticles, and the incorporation of drugs with different physical properties (ICG being amphiphilic and DOX being hydrophobic). These ICG–DOX nanoparticles have potential applications as drug delivery systems for combined chemotherapy and localized hyperthermia. Acknowledgments This work was supported by the Florida Department of Health Bankhead-Coley Cancer Research Program grant 08BB-11, and A.F.F was supported by NIH/NIGMS R25 GM061347. The authors would also like to acknowledge the technical help of Denny A. Carvajal, B.S. BME during the experimental phase of this research.

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