Thermosensitive, Near-Infrared-Labeled Nanoparticles for

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Thermosensitive, Near-Infrared-Labeled Nanoparticles for Topotecan Delivery to Tumors Elena V. Rosca,†,‡ Michael Wright,† Roman Gonitel,† Wladyslaw Gedroyc,§ Andrew D. Miller,† and Maya Thanou*,† †

Institute of Pharmaceutical Science, King’s College London, London, U.K. Department of Experimental Medicine, Imperial College London, London, U.K.

§

S Supporting Information *

ABSTRACT: Liposomal nanoparticles have proven to be versatile systems for drug delivery. However, the progress in clinic has been slower and less efficient than expected. This suggests a need for further development using carefully designed chemical components to improve usefulness under clinical conditions and maximize therapeutic effect. For cancer chemotherapy, PEGylated liposomes were the first nanomedicine to reach the market and have been used clinically for several years. Approaches toward targeted drug delivery using next generation “thermally triggered” nanoparticles are now in clinical trials. However, clinically tested thermosensitive liposomes (TSLs) lack the markers that allow tumor labeling and improved imaging for tissue specific applied hyperthermia. Here we describe the development of optically labeled TSLs for image guidance drug delivery and proof-of-concept results for their application in the treatment of murine xenograft tumors using the anticancer drug topotecan. These labeled TSLs also allow the simultaneous, real-time diagnostic imaging of nanoparticle biodistribution using a near-infrared (NIR; 750−950 nm) fluorophore coupled to a lipidic component of the lipid bilayer. When combined with multispectral fluorescence analysis, this allows for specific and high sensitivity tracking of the nanoparticles in vivo. The application of NIR fluorescence-labeled TSLs could have a transformative effect on future cancer chemotherapy. KEYWORDS: thermosensitive liposomes, topotecan, anticancer, focused ultrasound, multispectral fluorescence imaging

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INTRODUCTION Suboptimal delivery of cytotoxic chemotherapeutic agents is the main cause of serious side effects or failure for multiple cancer therapies.1 One approach to minimize toxicity is the use of a delivery protocol or vehicle that enhances tissue specificity, such as nanoparticles used to bind or encapsulate smallmolecule drugs.2,3 Owing to the enhanced permeability retention (EPR) effect, drug-bearing nanoparticles will accumulate into well vascularized tumors, resulting in an enhanced local concentration.4−6 This form of passive targeting means in principle that drugs can be effective at significantly lower dosages, minimizing undesirable side effects. 7,8 PEGylated liposomal drug−nanoparticles have been reported to operate in this way.9,10 For instance, Doxil (Caelyx) is a ∼90 nm particle with a polyethylene glycol (PEG) surface coat and lipid-encapsulated doxorubicin. This received EMA approval in June 1996, then FDA approval in June 1999, and was introduced into clinical practice to reduce doxorubicin cardiotoxicity.11 A drawback of such formulations is that the surface PEG layer improves in vivo stability of the nanoparticle but tends to actively delay the release of drug payload on reaching the target tissue.12,13 For this reason, we are interested in the development of “triggered” PEGylated nanoparticles for drug delivery in tumors. These triggered (or “stimuli © XXXX American Chemical Society

responsive”) nanoparticles are bimodal, in that they are designed for sufficient levels of stability in biological fluids under physiological conditions from the point of administration to the target tissue, whereupon they allow for the controlled release of the therapeutic agent on activation by changes in local endogenous conditions or through the application of an external/exogenous stimulus.14,15 Examples of triggered PEGylated drug−nanoparticles are now appearing.12,16 Triggerability has been explored using a variety of stimuli (pH, redox, enzymatic activation) and PEG modifications.17−20 However, the most advanced are the thermally triggered drug− nanoparticles that have progressed to clinical trials in recent years (Phase III HEAT Study in Primary Liver Cancer; Celsion.com).21,22 There are several types of thermosensitive liposomes developed and compared during the last two decades, however, ThermoDox, a commercial doxorubicinloaded version of the LTSL (low temperature sensitive liposomes), is probably the best developed.23,24 Due to their lower trigger temperatures (Tm), most TSLs deliver their drug Received: April 12, 2014 Revised: March 13, 2015 Accepted: March 31, 2015

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Figure 1. Spectral behavior of topotecan and XL750. (a) The excitation (solid lines) and emission (dotted) fluorescence spectra for topotecan vary significantly in buffers from pH 4.0 (red; inside liposomes) to pH 7.4 (blue; outside). Spectra were collected using the excitation/emission settings: pH 4.0 Em 536 nm, Ex 380 nm; pH 7.4 Em 536 nm, Ex 410 nm; instrument gain was the same for all scans. (b) The drug shows a mild fluorescence self-quenching at higher concentrations, but the pH sensitivity has a more significant contribution. (c) The increase in liposome-encapsulated topotecan fluorescence (at 640 nm emission maxima) before and after release by mild hyperthermia. (d) XL750-DSA shows strong near-infrared absorbance and fluorescence emission at 760 and 780 nm, respectively.

ability to image tumors and act as theranostics. These iTSLs are simultaneously formulated for triggered release of a therapeutic drug in response to externally induced hyperthermia, and equipped with a near-infrared fluorescence (NIRF) imaging component for real-time diagnostic imaging of the nanoparticle. The use of the NIR region of the spectrum (∼750−950 nm) minimizes problems with light absorption and scattering by body tissues (and particularly hemoglobin) allowing for improved transmission and excellent distinction from background signal, particularly when combined with multispectral analysis.38−42 These properties have enabled the use of optical imaging even in clinical trials.43−45 NIRF imaging still has significant limitations in terms of tissue penetration and occlusion, which has limited its wide adoption,46 however, in combination with suitable fluorescent agents it allows imaging up to 2 cm into the body, which is conducive to preclinical and even some clinical applications.47,48 Since the imaging technique is near real time in vivo NIRF can be used to investigate the biodistribution of labeled drugs or nanoparticles.49,50 In the case of triggerable nanoparticles, this information greatly contributes to the optimization of the trigger timing.

payload with the effects of mild hyperthermia thus reducing incidences of tissue overheating and related damage. In this case, external hyperthermia stimuli are indispensable for triggered release.25,26 The topic of optimum formulation and hyperthermia application has been the topic of recent research. Fast and slow releasing TSLs have been prepared and tested with mild hyperthermia. Although the fast doxorubicin releasing TSLs showed a lower Tm (and less stability in serum) than the slow doxorubicin release liposomes, they proved superior when combined with a two step mild hyperthermia protocol, suggesting that thermosensitivity and timings of hyperthermia applications are important for maximum antitumor effect.27 The hyperthermia stimulus may be realized in the clinic using radiofrequency (RF) ablation or high intensity focused ultrasound (HIFU or FUS) potentially in combination with an imaging technique, such as diagnostic ultrasound, X-ray computed tomography (CT), or magnetic resonance imaging (MRI).23,26,28−31 RF and HIFU techniques are in clinical use for tumor/tissue ablation and more recently to investigate the release of payloads from PEGylated drug− liposomes.32 The advent of nanoparticles bearing imaging agents (such as chelated gadolinium for MRI contrast enhancement)33−35 allows the targeted delivery of therapeutics combined with simultaneous real-time diagnostic imaging of nanoparticle biodistribution.36 The possibility of combining MRI contrast enhancement together with hyperthermia also provides the opportunity for real-time control and monitoring of the efficiency of TSLs mediated treatment.37 Here, we describe the early development of image-guided thermosensitive PEGylated liposomes (abbreviated as iTSLs) that have the

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RESULTS AND DISCUSSION Liposome Preparation and Characterization. In this study we prepare topotecan liposomes that have the imaging probes as part of the lipid bilayer (iTSLs). The selection of topotecan and the NIRF label XL750 has been based on their different emissive spectral characteristics and strong pHdependent variations in topotecan absorption and emission

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Molecular Pharmaceutics Scheme 1. Chemical Structures of Lipids and Drug Used To Prepare Labeled Thermosensitive Liposomes

A liposome composition was derived from ThermoDox (known to retain doxorubicin at 37 °C but rapidly release it at 42 °C) with the addition of two labeling lipids and modification of the loading parameters to suit topotecan. The lipid formulation was DPPC:MSPC:DSPE-PEG2000:Eu.DOTA.DSA :XL750.DSA at 80.2:9.7:5:5:0.1 (mol:mol), see Scheme 1 for the chemical structures of the components. Two of these lipids (Eu.DOTA.DSA, XL750.DSA) are synthetic conjugates, allowing the liposomes to be labeled respectively with europium and a commercially available NIR dye. The chosen synthetic routes were based on previous experience with the synthesis of Gd.DOTA.DSA.33−35,53 MSPC is introduced in the lipid composition to increase the instability at the Tm, enhancing the speed of drug release. This has been previously demonstrated in the ThermoDox formulation, which is the model used for this study.54 These resulting liposomes were loaded by incubation at 37.0 °C with the alkaloid anticancer drug topotecan, yielding labeled thermosensitive liposomes suitable for imaging (iTSLs). Representative data of iTSLs with and without loaded drug are shown in Table 1. These iTSLs were then characterized for their thermal triggerability and short/long-term stability in biological fluids.

characteristics. Figure 1 shows the spectral characteristics of the drug and the NIRF label. Topotecan fluorescence is highly pH dependent (Figure 1a), and while its emission shows selfquenching at high concentration (Figure 1b), the pH induced variation is more significant if the chosen excitation wavelength is in the region of high variation in the excitation spectra (about 390−480 nm). This is evident when topotecan labeled liposomes are excited at 410 nm (the peak for the neutral/ basic state). At this wavelength the liposomes show minimum fluorescence at 37 °C when topotecan is in the aqueous core (pH ∼ 4) and a great enhancement in intensity when heated at 45 °C for 2 min and drug is released in the solution (pH 7.4; Figure 1c). Figure 1d shows that XLA750 excitation and emission maxima are different to topotecan, allowing them to be easily distinguished using multispectral analysis. There is a concern that thermosensitive liposomal formulations may show some leakage of their content at physiological temperatures.51,52 Drug (e.g., doxorubicin) release from TSLs during their circulation in blood could lead to unwanted toxicity. To avoid this, optimization and testing of the formulation is required. C

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Molecular Pharmaceutics Table 1. Representative iTSL Characteristicsa concn

iTSL drug−iTSL

z-av size (diam.nm) 80.4 86.4

PDI

lipid (mg/mL)

topotecan (μg/mL)

0.23 0.23

4.3 ± 0.1 4.1 ± 0.1

70 ± 10

a

Sizing data collected by dynamic light scattering; lipid concentration estimated from Stewart assay; topotecan concentration measured by HPLC, n = 3 ± SD.

MSPC lipid was chosen as the main lysolipid to be tested in these iTSLs, as one of the most widely studied lysolipids.24 MSPC lysolipid however may not be the ideal lipid for TSLs. A study presented by Banno et al. indicated that the lipid may leave the formulation when the TSLs are in the blood circulation.55 The protocol used for topotecan loading takes advantage of the effects of different pH buffers inside and outside the liposomes. Topotecan shows changes in ionization state on acidification to pH ∼ 4. This affects the optical properties (absorbance and fluorescence excitation band peaks and intensities) and also the drug’s lipophilicity and hence retention by the liposome bilayer. Solubility is reduced by increasing acidity, but significant precipitation requires a pH below that used here (unlike doxorubicin). By formulating PEGylated liposomes with an internal aqueous core at pH 4.0 but using a physiologically compatible external buffer at pH of 7.4, the resulting pH gradient promotes drug entrapment within the liposome cavities on incubation near (but below) the Tm.56,57 Since the lipid membrane porosity is thermally sensitive, warming the purified iTSLs to above the Tm allows the trapped drug to be released. This release can be easily observed as a function of fluorescence intensity changes with time (Figure 2). The increase in fluorescence intensity with drug release is sufficiently distinct (Figure 1c and Figure 2) that it provides a ready biophysical means to study the thermally triggered release of topotecan from iTSLs as a function of differing incubation temperature and time. Thermally triggered release from iTSLs was demonstrated using a Maestro EX multispectral imager (with blue/green excitation/emission filters) using physiologically equivalent (37 °C, 3 min) or hyperthermic samples. The hyperthermic samples showed significantly higher fluorescence, with a spectral profile that matched that expected for topotecan (see Figure S4 in the Supporting Information). Thereafter, the effectiveness of the thermal trigger was assessed by quantifying the increase in topotecan fluorescence as a function of incubation time at different incubation temperatures and buffers (see Figure 2). At 37 °C, in HEPES buffer iTSLs appeared stable and minimal drug release was observed in the short term (10 min). As the temperature was increased to 41 °C, the release rate increased and was found to be complete within 8−9 min (Figure 2a). The stability in different biologically relevant media was then studied in terms of drug release taking place over 45 min of incubation at 37 °C (Figure 2b). The results suggest adequate serum or plasma stability to about 45 min at 37 °C. Incubation for significantly longer time (9 h) at the same temperature gave a 40% drug release in HEPES buffer. The release of topotecan in 50% plasma starts earlier by reaching 50% drug release at 3 h. This has been previously observed in TSL formulations, and similar data on long-term stability have been recently published; doxorubicin containing LTSL

Figure 2. Temperature responsiveness and serum/plasma stability. (a) Topotecan release from iTSLs as a function of time and incubation temperature, as assessed by the increase in fluorescence. (b) The integrity of TSLs and the short-term effects on drug release of incubation at 37 °C in HEPES buffer (pH 7.4) or in the presence of 30% or 50% fetal calf serum, and 50% human plasma. (c) Demonstration of the longer term drug release on incubation at 37 °C for up to 20 h in buffer and 50% human plasma. Data points shown are an average of three replicates, and the error bars are ± SD.

(ThermoDox) long-term stability at 37 °C studies in 50% plasma suggested complete drug release within 3 h.58 The small difference in the relative stability at 3 h in 50% plasma may be explained by the inclusion of additional imaging lipids (Eu.DOTA.DSA 5 mol % and XL750.DSA 0.1 mol %). In Vivo iTSLs Kinetics and Mediated Tumor Growth Suppression. Following these experiments, iTSLs were administered by iv injection and nanoparticle kinetics were then analyzed as a function of time. NIRF-iTSLs can be imaged clearly in circulation from about 45 min postinjection as the nanoparticles reach and collect within the two flank tumors D

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Figure 3. In vivo particle kinetics in tumors generated with IGROV-1 cells subcutaneously implanted to the flanks of SHO mice. (a) Illustration of the highlighted vasculature due to iTSLs presence 45 min after injection. (b) Illustration of the accumulation of iTSLs in tumors over time, in one representative animal. The images were acquired using the same NIR acquisition setting at each time point; images were unmixed using the Maestro software, quantified using pixel intensity, and false colored using ImageJ. (c) Average signal intensity corresponding to the particle kinetics over time using the quantified images of each acquisition time point. The measurements are an average of 8 tumors with the error bars ± SD.

uptake by the whole immersed part as indicated by bioimaging, and it was suspected that hyperthermia had affected the surrounding tissues. Tissue temperatures were monitored throughout using small probe thermocouples placed at the base of the tumor away from the heating source (see Figure S5 in the Supporting Information). An apparent increase in drug fluorescence signal was detected within 5 min from the beginning of heating (Figure 4a). This is considered to correspond with heat-triggered drug release into tumor environments in accordance with iTSL biophysical drug release data (Figure 2). The observation was also found to correlate agreeably with enhanced reductions in the growth of iTSLloaded and hyperthermia-treated tumors (heated; 42 °C) in comparison to the control situation where heat treatment was withheld (nonheated) (Figure 4b). In order to characterize the effects of tumor heat treatment vs non heat treatment in more detail, confocal microscopy of iTSL tumor sections was performed (Figures 4c,d). Postheating, fluorescence distribution was found to be more diffuse within this tumor section, a result consistent with triggered drug release and intratumoral diffusion of free topotecan (Figure 4d). Further in vivo mechanistic studies were performed using “pseudotumors”. GelTrex samples were loaded with iTSLs at dose levels equivalent to 5% of the normal total iv-injected dose. Using optical imaging (Figure 3), we prior estimated that signal-to-background noise ratios were adequate for clear observation when at least 5% of the initial iv-injected dose of nanoparticles partitioned into tumor volumes 1 h postinjection. These iTSL-loaded Geltrex “pseudotumors” were then implanted into mice and imaged. NIR multispectral fluorescence imaging was carried out to characterize the presence of iTSLs and drug in the absence of heat treatment and after 5 min of heat treatment to 42 °C (Figure 5 right pseudotumor). This demonstrated an increase in drug fluorescence signal over nonheated control (Figure 5c) 5 min after heating iTSL pseudotumor (Figure 5d). The topotecan signal coming from

(Figure 3a). This accumulation of iTSLs was monitored over time and shown to promote a distinct increase in the XL750.DSA fluorescence signal, using the following instrument settings: excitation band pass 704 nm (684 to 729 nm effectively) and emission long pass 745 nm, and images acquired over 740 to 950 nm in 10 nm steps (see Figure 3b and Supporting Information Figure S4). The signal apparent in the neck/head of the animal is considered to be due to the presence of the iTSLs in the vasculature, in particular the subcutaneous vessels, and not due to uptake to the brain. The maximum signal occurs 15−30 min postinjection and decays rapidly over the following hours, which supports this hypothesis. Conversely, the NIRF signal in the liver and spleen increases over time and has been confirmed by ventral imaging. Semiquantification of images based on the fluorescence intensities of the tumor as a function of time provides the illustrated nanoparticle kinetic profile (Figure 3c). These kinetics are based on the intensity of the NIRF signal coming from the liposomes. The data indicate that signal is gradually increasing during the first 5 h while minimum changes at the fluorescent signal appear from 5 to 24 h. These results were obtained from real-time imaging of nanoparticle tumor accumulation. Detecting the level on nanoparticles in tumors in real time is an important factor for the success of image guided triggered drug delivery. In situ drug release from tumor-localized iTSLs was observed upon warming iTSL-loaded tumors with an external heat source (water circulating copper heat block at 45 °C; this water temperature was chosen as it gave 42 °C as measured by the thermocouples; Supporting Information Figure S5). The method was developed to allow a local heating effect (only in the tumors), and although the heating is superficial in comparison to the heating via FUS, in the case of the subcutaneous tumors used in this study the heating was homogeneous as demonstrated in the Supporting Information (Figure S5). Previous attempts of using mice immersed in a thermostated water bath resulted in increased nanoparticle E

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the implanted Geltrex pseudotumors is more evident as these were implanted under the skin (reducing tissue occlusion), and avoids the hemoglobin absorbance that would be expected from the blood vessels of a real tumor. In these pseudotumors the pH was buffered allowing the topotecan signal to be imaged, however in real tumor cells the pH may be slightly more acidic (pH ∼7.00) compared to the physiological pH of normal tissues (pH ∼7.5).59,60 This could have an effect on topotecan’s fluorescent emission. The principle of this approach has been used in the measurement of the pH within tumors or to quantify fluorescence.61 Quantification of iTSL partition from circulation into the tumors was determined using fluorescence signals from both NIR and topotecan (in vivo) and europium labeled lipids incorporated into iTSLs. Europium was assessed in excised tumors using a postlabeling assay.62 Using an assessment of this second lipid provides a validation of the biodistribution monitoring. Based upon the near linear concentration dependence of fluorescence at low concentrations, a standard curve was generated from sequences of images of Geltrex pseudotumors loaded with known concentrations of iTSLs. Eu-fluorescent calibration was prepared in lysed excised untreated tumors. This was then used to calibrate fluorescence signal intensities (NIR and Eu) from iTSL-loaded tumors. Results from both fluorescent labels were consistent with a tumor accumulation of about 15% of the injected dose (Figure 6d). Furthermore, when fluorescence coregistration was performed with iTSL-loaded tumors in vivo 8 h post iv injection, color merging of the NIR signal (red) and drug signal (green) indicated colocalization of both (Figure 6a−c). These results are consistent with colocalization of both drug and lipid into the tumor. The need for more advanced nanoparticle-delivery technologies to improve drug pharmacokinetics and pharmacodynamics is well-known. Here we investigate the construction of iTSLs for the heat-triggered release of drug locally in response to external stimuli, which are also NIR-labeled for real-time observation of iTSL delivery using fluorescence imaging. Our results indicate that the reported iTSLs do release their payload in response to a thermal trigger (Figures 2, 4−6), and they can be seen to partition from circulation to tumor by observation of the NIR fluorescent label (Figure 3), and apparently to release drug in vivo following heat treatment (Figures 4 and 6). It has been suggested that the use of MRI-guided focused ultrasound (MRgFUS) may be effective for the triggered release of doxorubicin in tumors from thermosensitive nanoparticles.23,51,63 Indeed, MRgFUS systems have been adapted for preclinical investigations in rodents64 although current demonstrations of actual MRI-guided FUS-triggered release of drug in tumors in the preclinical setting are limited to dose painting experiments, involving the coencapsulation and release of the drug along with small molecule MRI contrast agents.65−67 Using NIR optical imaging we were able to demonstrate the time parameters required for significant tumor accumulation of NIR-labeled iTSLs, a key result needed in order to ensure sufficient accumulation into tumor prior to the use of hyperthermic trigger, in order to maximize the efficacy against tumors (see Scheme 2). In this trial study, the hyperthermia effect was induced using a heating block intended to induce controlled, local tumor heating (Figure S5 in the Supporting Information), making use of a short heating time (5 min vs ∼1 h which is currently more typical68). Other reported hyperthermia treatments have made

Figure 4. Drug delivery and efficacy upon heating (hyperthermia at 42 °C). (a) Quantification of the fluorescence signal generated by topotecan before and after heating (at 4 h postinjection). Images were acquired by multispectral fluorescence imaging before and after heating and unmixed by subtracting the background fluorescence, and the intensity in the tumor was measured by counting the intensity of the pixels. (b) Inhibition of tumor growth over time post iTSLs accumulation and with or without the application of heat (N = 4). (c, d) Sections representing a typical drug distribution in a nonheated and in a heated tumor (magnification 100×) of tumor sections (z-stack) imaged by confocal microscopy (sections representative of 3 processed and imaged tumors per group heated/nonheated). The topotecan gives the green fluorescence while the nucleus of the cells is counterstained by DAPI (blue).

Figure 5. In vivo drug release confirmation by fluorescence monitoring. The mouse was implanted with GelTrex “pseudotumors” mixed with TSLs (5% of normal iv injected dose). The upper panels show the NIR signal before (a) and after (b) heating the right-hand pseudotumor. The lower panels show drug fluorescence before (c) and after (d) heating, thereby demonstrating an approximately 2-fold increase in fluorescence in heated pseudotumor versus nonheated control.

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Figure 6. Tumor accumulation of iTSLs in non thermally treated tumors. Illustration of (a) the in vivo accumulation of the particles using the NIR signal which indicates the presence of XL750-DSA lipid and (b) topotecan based on its fluorescence. (c) Colocalization of the two signals in tumors 8 h post iv injection. (d) Quantification of iTSLs accumulated in the tumor in vivo as detected by NIR fluorescent signal analysis or ex vivo (after excision and lysis) by europium (Eu) fluorescent signal analysis.

Scheme 2. Schematic of Thermally Triggered Theranostic Nanoparticle Methods in Vivoa

a iTSLs are injected iv and allowed to accumulate in the tumor due to the EPR effect, while being monitored by NIR-fluorescence imaging. Once the concentration is thought optimal, the therapeutic agent is locally released with brief (∼5 min) hyperthermia. Exposure of the cancer cells to a high local concentration of agent within a short period of time maximizes cytotoxicity.

use of hot water whole leg bath immersion protocols, but these appear to induce increased general permeability of tissues to nanoparticle incursions other than just tumor volume.21,58 The obvious questions remaining are when would be the optimal time points in which to apply heat treatment, for how long, until when, and in what time intervals. Optimized selective drug delivery is dependent on several parameters but, in particular, the nanoparticle pharmacokinetics (particokinetics), the rate of accumulation of drug−nanoparticles into tumors (biodistribution), and the stability of these nanoparticles once at the site of action.69 Different studies with TSLs initiate drug release either immediately or 24 h postadministration of the liposomes, dependent on the design release profile (fast or slow release27).30 However, the optimal time (thermal dosing) of hyperthermia application for maximum drug delivery remains undefined. From most recent studies presented it is evident (and well accepted) that local hyperthermia can substantially increase the tumor’s vascular permeability and the local drug/ nanoparticle distribution. It is also clear that hyperthermia application needs to be combined with image guidance, and the combination may provide a valuable tool in treating cancers. Overall we have demonstrated the formulation of theranostic PEGylated nanoparticles that are loaded with an anticancer drug topotecan, possess the property of thermal triggerability, and are also labeled with a NIR fluorescent dye for simultaneous, near real time imaging of nanoparticle biodistribution and drug pharmacokinetics in vivo. We believe that this demonstrates proof of principle for image-guided, nano-

particle-mediated anticancer therapy, an important step forward toward personalized therapies.

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MATERIALS AND METHODS Materials and General Methods. 1,2-Dipalmitoyl-snglycero-3-phosphocholine (DPPC; 16:0 PC), 1-stearoyl-snglycero-3-phosphocholine (MSPC; 18:0 Lyso PC), and (ωmethoxy-polyethylene glycol 2000)-N-carboxy-1,2-distearoylsn-glycero-3-phosphoethanolamine (PEG2000-DSPE) were purchased from Avanti Polar Lipids (Alabaster, AL, USA). DOTANHS-ester was purchased from Macrocyclics (Dallas, TX, USA) and XenoLight750-NHS-ester from PerkinElmer (Waltham, MA, USA). Cell media were from Life Technologies (Carlsbad, CA, USA) while other materials were from SigmaAldrich (St. Louis, MO, USA) and were of analytical grade. Other lipids were synthesized as described below. 1H (400 MHz) and 13C (100 MHz) NMR spectra were recorded on a Bruker Avance 400 spectrometer using residual chloroform or dichloromethane as internal standards. Results are reported as chemical shifts in ppm from TMS, with peaks described as s = singlet, br = broad singlet, d = doublet, t = triplet, q = quartet, m = multiplet, and coupling constants J given in hertz (Hz). Mass spectroscopy was carried out on Thermo LCQ DECA XP or Agilent HP1100 MSD spectrometers depending on availability. Analytical HPLC was carried out using an Agilent 1100 series instrument equipped with a multiwavelength diode array detector, a 1260 Infinity fluorescence detector, a Polymer Laboratories PL-ELS-2100 evaporative light scattering detector, G

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Molecular Pharmaceutics and a 5 cm Hypersil C18 5 μm reverse-phase column. Synthesized lipids were analyzed using the gradient 0 min, 100% water, 2.5 mL/min; 1 min, 100% water; 11 min, 100% ACN; 11 min, 100% ACN; 23 min, 100% methanol; 25 min, 100% methanol; 27 min, 100% water, 1.8 mL/min; 30 min, 100% water, 2.5 mL/min; and they showed purity at least 95% (see Figure S6 in the Supporting Information). Thin layer chromatography (TLC) was carried out on F254 silica gel 60 plates, with spots visualized by UV illumination or vanillin/ ninhydrin staining and developed with a heat gun. Flash column chromatography was performed on 40−63 μm silica gel. Synthesis of Lipids. N,N-Distearylamidomethylamine (DSA). DSA was synthesized according to Kamaly et al.33. 1H (400 MHz; CD2Cl2; 296 K): δ 3.84 (s, 2H, OCCH2NH2), 3.29 (t, J = 8.0 Hz, 2H, OCNCH2), 3.11 (t, J = 7.8 Hz, 2H, OCNCH2), 1.50 (m, 4H, OCNCH2CH2), 1.25 (s, 60H, alky chain CH2), 0.88 (t, J = 6.3 Hz, 6H, CH3). 13C (100 MHz; CD2Cl2; 296 K): δ 166.6 (OCN) 48.8 and 48.1 (OCNCH2), 41.6 (OCCH2NH2), 31.1 (CH3CH2CH2), 30.9−30.8 (alkyl chain CH 2 ), 29.9 (OCNCH 2 CH 2 CH 2 CH 2 ), 28.8−28.3 (OCNCH2CH2CH2), 24.2 (CH3CH2), 15.4 (CH3). TLC (15% MeOH in CH2Cl2 with 0.5% NH3) gave Rf 0.55 with the DSA spot showing red after sequential vanillin and ninhydrin stains. HPLC tR = 13.3 min (see Supporting Information Figure S3). ESI-MS: [M + H]+ 579.7 m/z (expect 578.6 m/z for C38H78N2O). N′-XenoLight750-N,N-distearylamidomethylamine (XL750.DSA). DSA (4.2 mg; 7.3 μmol; 1) was dissolved under nitrogen in dry DCM (0.2 mL) with distilled triethylamine (20 μL, 0.14 mmol). XenoLight750-NHS (1 μmol) dissolved in dry DMSO (100 μL; requires vigorous vortexing) was added, the flask protected from light, and the mixture gently stirred. TLC (15% methanol in CH2Cl2 with 0.5% NH3) showed conversion of separate DSA (Rf 0.55) and XenoLight750-NHS (Rf 0.10) spots to a streak (Rf 0.40−0.65) over 5 h. The reaction was then stopped and the reaction mixture dried in vacuo before purification by flash column chromatography (2 mL) loaded in CH2Cl2 and eluted with 5% (DSA; colorless); 15−20% (conjugate; blue; Rf 0.65); then 30% (side product; colorless; Rf 0.45−0.55) MeOH with unconjugated dye retained on the column. The XL750-DSA fractions were combined and dried in vacuo to give a dark-blue solid with an estimated 70% yield. This was dissolved to 1 mg/mL in chloroform and then stored at −20 °C. 1H (400 MHz; CD2Cl2/CD3OD; 296 K): δ 7.62 (m, 3H, dye), 7.48 (m, 3H, dye), 4.11 (m, 5H, dye), 3.85 (m, 2H, OCCH2NH2), 3.33 (s, 2H, OCNCH2), 3.26 (s, 3H, dye), 3.12 (s, 2H, OCNCH2), 3.00 (q, 4H, dye), 2.04 (s, 3H, dye), 1.59 (m, 4H, OCNCH2CH2), 1.27−1.18 (m, 60H, alkyl chain CH2), 0.82 (t, J = 7.5 Hz, 6H, CH3). 13C (100 MHz; CD2Cl2; 296 K): δ 131.3 (dye), 129.1 (dye), 70.8 (dye), 68.4 (dye), 46.4 (OCNCH2), 39.2 (OCCH2NH2), 34.7 (dye), 30.7−29.3 (alky chain CH2), 24.1 (dye), 23.4 (CH3CH2), 14.2 (CH3), 11.1 (dye), 8.8 (dye). The structure of the XenoLight750 dye is unavailable so assignment information is limited. HPLC: tR = 26.7 min (see Supporting Information Figure S3); λabs/λem (CHCl3) 755/775 nm. Europium(III) 2-{4,7-bis-carboxymethyl-10-[(N,N-distearylamidomethyl-N′-amidomethyl]-1,4,7,10-tetraazacyclododec-1yl}acetic acid (Eu.DOTA.DSA) was synthesized by adaption of the protocol of Kamaly et al.33 In brief, DOTA-NHS-ester (100 mg, 0.120 mmol) and DSA (80.2 mg, 0.139 mmol) were dissolved in dry CH2Cl2 (40 mL). Distilled Et3N (67 μL, 0.48

mmol) was added and the mixture stirred under N2 for 12 h at 35 °C. The solution was dried in vacuo and purified by flash chromatography loaded in a 10% CH 2 Cl 2 :MeOH:NH 2 (34.5:9:1) mixture in CH2Cl2 and eluted with increasing concentration to 100% of the solvent mixture. Fractions containing the target were identified by HPLC, combined, and dried to give a white hydroscopic solid (57.0 mg; 49%). 1H (400 MHz; CD2Cl2/CD3OD): δ 3.45 (br, 2H, NCH2CONH), 3.10 (br, 6H, NCH2COOH), 3.00 (br, 2H, OCNCH2), 2.80 (br, 16H, NCH2CH2N), 2.28 (br, 2H, OCNCH2), 2.16 (br, 2H, OCCH2NH), 1.44 (m, 4H, OCNCH2CH2), 1.18 (s, 60H, alky chain CH2), 0.80 (t, J = 6.6 Hz, 6H, CH3). 13C (100 MHz; CD2Cl2/CD3OD; 296 K): δ 47.0 (OCNCH2), 41.5−38.5 (NCH2CH2N and NCH2COOH), 32.3 (CH3CH2CH2), 31.5− 28.5 (alkyl chain CH2), 22.9 (CH3CH2), 14.1 (CH3); others could not be distinguished. HPLC: tR = 21.2 min (see Supporting Information Figure S3). ESI-MS [M + H] + gave 965.7 m/z (expected 964.8 for C54H104N6O8) with major fragments seen at 579.6, 522.3, 444.1, and 387.1 m/z corresponding to DSA 1, (C18)2NH, DOTA-glycine, and DOTA respectively. Europium complexation was effected by suspension of DOTA-DSA (25.2 mg, 0.026 mmol) in a vigorously stirred aqueous solution (5 mL) of europium(III) chloride (10.5 mg, 0.029 mmol of hydrate) heated at 90 °C for 12 h under N2. After settling, the excess water was removed and minimal CH2Cl2 added to dissolve the lipid complex. After vigorous mixing with equal amounts of deionized water, the emulsion was separated by centrifugation and the CH2Cl2 layer collected and dried in vacuo to give a white power (28 mg; 96%). HPLC: tR = 21.4 min. ESI-MS [M + H] + gave 1115.7 m/z or [M − H]− 1113.6 m/z (expected 1114.7 for C54EuH101N6O8) with major fragments seen at 1070.5, 1026.6, 982.6 m/z corresponding to loss of COO, 2 × COO, and 3 × COO respectively. Preparation of iTSLs with Topotecan. All lipids were stored in aliquots of 10 mg/mL in either CHCl3 or MeOH/ CHCl3 50:50 v/v. 80.2 mol % DPPC, 9.7 mol % MSPC, 5 mol % -PEG2000-DSPE, 5 mol % Eu.DOTA.DSA, and 0.1 mol % XL750.DSA were combined in a round-bottom flask to give an appropriate total mass of lipid (5−10 mg). The solvent was slowly evaporated in vacuo to ensure a thin and even film formation. This was hydrated in 1 mL of 300 mM ammonium sulfate, at pH 4.0, and treated 5 times by freeze/thaw in liquid nitrogen and hot water to fragment the film. The resulting suspension was then sonicated at 60 °C for 10 min forming a milky blue/white liquid, before extrusion at least 3 times through a 100 nm membrane using a Northern Lipids (Burnaby, Canada) LIPEX extruder heated to 55 °C and pressurized to about 10−20 bar. The external buffer was then exchanged to sterile 20 mM HEPES pH 7.4 with 5% glucose (w/v) using a PD10 size exclusion column (Amersham, Buckinghamshire, U.K.). The resulting slightly cloudy blue suspension was sized using Nanoseries Nano ZS (Malvern Instruments, Worcestershire, U.K.) before incubation with topotecan hydrochloride (0.25 mM aq) (Sigma-Aldrich, St. Louis, MO, USA) at 37 °C for 2 h.70 This step used a Thermocycler (Mastercycler Personal, Eppendorf, Stevenage, U.K.) to accurately control the temperature. The excess drug was removed using a PD10 column loaded with HEPES buffer, giving a clear blue/green suspension. The size of the liposomes was recorded, samples (50−100 μL) were taken to quantify the H

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maintained according to the regulations of Home Office, with food and water available at libitum. Mice were inoculated with IGROV-1 (ovarian cancer) cells subcutaneously in the back flank, 3 × 106 cells suspended in complete media and mixed 1:1 with Geltrex (Life Science, U.K.). Tumors were formed at about 14 days post inoculation. When the tumor reached a volume of approximately 100 mm3 (measured by calipers), the treatment commenced. The mice were placed under anesthesia using isoflurane gas, nanoparticles were injected iv, and the distribution was monitored using an in vivo multispectral imaging system (Maestro EX, PerkinElmer, Waltham, MA, USA). For the first hour images were collected every 10−15 min and then every hour subsequently. Nanoparticle distribution images were acquired using the NIR filters set to track XL750-DSA, using the same exposures between the time points. The fluorescence of the topotecan was also monitored using the blue (445 nm)/green (503−555 nm) excitation/ emission filter combination.72,73 After 4 h the tumor was heated to hyperthermia (42−43 °C) for 5 min using an external heating source (1 cm2 copper block heated by circulating water). Localized tumor temperature was monitored with a fine tip T-type thermocouple, and images pre- and postheating were acquired. Mice (N = 3) were humanely euthanized (immediately after hyperthermia or without hyperthermia) before tumors and organs were collected and imaged ex vivo and the tumors frozen and/or preserved in 1% formalin. Other treated mice (n = 4) were maintained for 3 weeks, and the tumor growth was monitored by caliper measurements. At the termination of the experiment, all mice were humanely euthanized. Europium was detected in excised and lysed tumors using the DELPHIA time resolved fluorescence assay (PerkinElmer, U.K.). Analysis of Tumor Sections Using Fluorescent Microscopy. Fixed tumors were sliced on a cryostat and mounted on slides. These were incubated with a PBS solution containing DAPI (50 μg/mL) for nuclear counterstaining. Images were collected using a Nikon confocal microscope (A100, Amsterdam, The Netherlands) at 20× and 100× magnification. The z-stack sequence was collected to image a tissue volume of dimension 212 × 212 × 63.5 μm. NIR Imaging and Statistical Analysis. Images were quantified by first unmixing the NIR signal from the background. This was accomplished using the Maestro EX software (PerkinElmer). The tumor was then outlined, and the fluorescence intensity was measured via pixel counts in the region of interest. The injected dose was calculated from the analysis of images of subcutaneous injection of Geltrex mixed with known amounts of liposomes. The in vivo release of the topotecan was also modeled using mock tumors of Geltrex loaded with known amount of liposome that were injected at the back flank of mice and treated with hyperthermia. The temperature was monitored in vivo with a fine tip thermocouple placed on the surface of tumor, away from the heating source. Several points around the tumors were checked to ensure that the entire tumor tissue had reached a homogeneous temperature (Supporting Information Figure S5). The statistical analysis was done using GraphPad (GraphPad Software, Inc., La Jolla, CA, USA) with ANOVA analysis used to determine statistical significance. The kinetic curve was checked for significance (increase of dose over time) using a post-test for linear trends.

lipid and topotecan concentration, and the rest was stored at 4 °C. The lipid was quantified using a modified version of the Stewart assay.71 In brief, 50 μL of liposomes was mixed with 150 μL of water and 200 μL of 1:1 MeOH:CHCl3 and then vortexed vigorously to an emulsion. The sample was centrifuged (4000g; 2 min) to fully separate the solvent layers. An aliquot (70 μL) of the organic layer was combined with Stewart reagent (5 μL, FeCl3/NH4SCN aq), and the combination was vortex mixed again and then centrifuged. An aliquot (50 μL) was then transferred to a glass 96-well plate (Cayman Chemical, Ann Arbor, MI, USA) and the absorbance at 455 nm measured on a plate reader (Infinite 200 Pro, Tecan, Männedorf, Switzerland) compared to known standards. The drug concentration was measured by HPLC using a 30 × 4.6 mm Hypersil C18 5 μm reverse-phase column, initially loaded in deionized water containing 0.1% trifluoracetic acid (TFA) before elution at 3.5 mL/min using the gradient 0 min 0% MeCN; 2 min 0%; 4 min 50%; 5 min 50%; 6 min 0%; 7 min 0%. Drug was detected by evaporative light scattering (ELS) at 3.6 min and by the acid-state absorption band (380 nm) at 3.5 min (see Figures S1 and S2 in the Supporting Information). Drug amounts were quantified directly from 50 to 100 μL liposome samples after calibration against known standards. Thermally Triggered Drug Release. The thermal trigger of drug release from iTSLs was assessed by fluorescence. Topotecan has a pH-sensitive UV/visible absorbance profile that shifts from an absorbance of 385 nm at pH ∼ 6.5 to 414 nm at pH > 7.5. This results in topotecan solutions changing from colorless in acidic buffers to yellow in neutral/basic ones. Since the liposome formulation contains a low internal pH, we take advantage of this shift in absorbance maximum to greatly enhance the difference in topotecan fluorescence on release from liposomal encapsulation. This is valuable since, unlike doxorubicin, topotecan’s fluorescence self-quenching on liposomal encapsulation is relatively weak. By measuring this change in fluorescence intensity on incubation at various temperatures, we can quantify the amount of released drug. Studies were carried out with separate samples (50 mL) of 1:20 diluted liposomes, incubated in a Thermocycler (37−41 °C; 0− 15 min) before cooling to room temperature, transferring to a 96-well plate, and measuring their absorbance (414 nm; bandwidth 9 nm) and fluorescence (Ex 414 nm bandwidth 9 nm; Em 530 nm bandwidth 20 nm) using the Tecan plate reader. iTSL Stability in Biological Buffers. The long-term stability of our iTSL formulations was evaluated in buffer (20 mM HEPES with 5% glucose, pH 7.4) and then buffer containing 30−95% fetal bovine serum (FBS) or 50% human plasma. The study was carried out as described in the previous section, only that the dilution was into the appropriate buffer and incubation was at 37 °C for up to 24 h. Fluorescence readings were taken using the Tecan plate reader as previously described. At the end of the study, the samples were heated to 43 °C for 3 min to confirm the retained capability of drug release despite prolonged incubation in serum and/or plasma. In Vivo iTSL Kinetics and Mediated Tumor Growth Suppression. All procedures on animals were conducted in accordance with UK Home Office regulations and the Guidance for the Operation of Animals (Scientific Procedures) Act (1986). Female, 4−6 week old SHO mice (SCID hairless mice) were purchased from Charles River (Wilmington, MA, USA) and I

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multidrug resistance in adriamycin resistant ovarian cancer cells (NCI/ ADR-RES). Int. J. Pharm. 2012, 431 (1−2), 222−9. (9) Maruyama, K. Intracellular targeting delivery of liposomal drugs to solid tumors based on EPR effects. Adv. Drug Delivery Rev. 2011, 63 (3), 161−9. (10) Gabizon, A.; Shmeeda, H.; Horowitz, A. T.; Zalipsky, S. Tumor cell targeting of liposome-entrapped drugs with phospholipidanchored folic acid-PEG conjugates. Adv. Drug Delivery Rev. 2004, 56 (8), 1177−92. (11) Gabizon, A.; Shmeeda, H.; Barenholz, Y. Pharmacokinetics of pegylated liposomal Doxorubicin: review of animal and human studies. Clin. Pharmacokinet. 2003, 42 (5), 419−36. (12) Torchilin, V. Multifunctional and stimuli-sensitive pharmaceutical nanocarriers. Eur. J. Pharm. Biopharm. 2009, 71 (3), 431−44. (13) Andresen, T. L.; Jensen, S. S.; Jorgensen, K. Advanced strategies in liposomal cancer therapy: problems and prospects of active and tumor specific drug release. Prog. Lipid Res. 2005, 44 (1), 68−97. (14) Miller, A. D. Lipid-based nanoparticles in cancer diagnosis and therapy. J. Drug Delivery 2013, 2013, 165981. (15) Miller, A. D. Delivery of RNAi therapeutics: work in progress. Expert Rev. Med. Devices 2013, 10 (6), 781−811. (16) Romberg, B.; Hennink, W. E.; Storm, G. Sheddable coatings for long-circulating nanoparticles. Pharm. Res. 2008, 25 (1), 55−71. (17) Carmona, S.; Jorgensen, M. R.; Kolli, S.; Crowther, C.; Salazar, F. H.; Marion, P. L.; Fujino, M.; Natori, Y.; Thanou, M.; Arbuthnot, P.; Miller, A. D. Controlling HBV replication in vivo by intravenous administration of triggered PEGylated siRNA-nanoparticles. Mol. Pharmaceutics 2009, 6 (3), 706−17. (18) Drake, C. R.; Aissaoui, A.; Argyros, O.; Serginson, J. M.; Monnery, B. D.; Thanou, M.; Steinke, J. H.; Miller, A. D. Bioresponsive small molecule polyamines as noncytotoxic alternative to polyethylenimine. Mol. Pharmaceutics 2010, 7 (6), 2040−55. (19) Kolli, S.; Wong, S. P.; Harbottle, R.; Johnston, B.; Thanou, M.; Miller, A. D. pH-triggered nanoparticle mediated delivery of siRNA to liver cells in vitro and in vivo. Bioconjugate Chem. 2013, 24 (3), 314− 32. (20) Yingyuad, P.; Mevel, M.; Prata, C.; Furegati, S.; Kontogiorgis, C.; Thanou, M.; Miller, A. D. Enzyme-triggered PEGylated pDNAnanoparticles for controlled release of pDNA in tumors. Bioconjugate Chem. 2013, 24 (3), 343−62. (21) Kong, G.; Anyarambhatla, G.; Petros, W. P.; Braun, R. D.; Colvin, O. M.; Needham, D.; Dewhirst, M. W. Efficacy of liposomes and hyperthermia in a human tumor xenograft model: importance of triggered drug release. Cancer Res. 2000, 60 (24), 6950−7. (22) Kong, G.; Braun, R. D.; Dewhirst, M. W. Hyperthermia enables tumor-specific nanoparticle delivery: effect of particle size. Cancer Res. 2000, 60 (16), 4440−5. (23) de Smet, M.; Langereis, S.; van den Bosch, S.; Grull, H. Temperature-sensitive liposomes for doxorubicin delivery under MRI guidance. J. Controlled Release 2010, 143 (1), 120−7. (24) Landon, C. D.; Park, J. Y.; Needham, D.; Dewhirst, M. W. Nanoscale Drug Delivery and Hyperthermia: The Materials Design and Preclinical and Clinical Testing of Low Temperature-Sensitive Liposomes Used in Combination with Mild Hyperthermia in the Treatment of Local Cancer. Open Nanomed. J. 2011, 3, 38−64. (25) Ta, T.; Porter, T. M. Thermosensitive liposomes for localized delivery and triggered release of chemotherapy. J. Controlled Release 2013, 169 (1−2), 112−25. (26) Grull, H.; Langereis, S. Hyperthermia-triggered drug delivery from temperature-sensitive liposomes using MRI-guided high intensity focused ultrasound. J. Controlled Release 2012, 161 (2), 317−27. (27) Li, L.; ten Hagen, T. L.; Haeri, A.; Soullie, T.; Scholten, C.; Seynhaeve, A. L.; Eggermont, A. M.; Koning, G. A. A novel two-step mild hyperthermia for advanced liposomal chemotherapy. J. Controlled Release 2014, 174, 202−8. (28) Negussie, A. H.; Yarmolenko, P. S.; Partanen, A.; Ranjan, A.; Jacobs, G.; Woods, D.; Bryant, H.; Thomasson, D.; Dewhirst, M. W.; Wood, B. J.; Dreher, M. R. Formulation and characterisation of magnetic resonance imageable thermally sensitive liposomes for use

* Supporting Information S

HPLC analysis of topotecan and lipids, fluorescence imaging of topotecan bearing liposomes, and temperature recording of thermocouple during tumor heating in vivo. This material is available free of charge via the Internet at http://pubs.acs.org. Corresponding Author

*Institute of Pharmaceutical Science, King’s College London, Britannia House, 7 Trinity Street, London, SE1 1DB, U.K. Tel: +44 (0) 20 7848 7530. E-mail: [email protected]. Present Address

‡ E.V.R.: Department of Biological Sciences, University of Hull, Hull, HU6 7RX, U.K.

Notes

The authors declare no competing financial interest.

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ACKNOWLEDGMENTS All authors acknowledge the financial support of EPSRC (Engineering and Physical Sciences Research Council, U.K.) that funded this project.

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ABBREVIATIONS USED CT, computed tomography; DPPC, 1,2-dipalmitoyl-sn-glycero3-phosphocholine; DSA, N,N-distearylamidomethylamine; DSPE-PEG2000, (ω-methoxy-polyethylene glycol 2000)-Ncarboxy-1,2-distearoyl-sn-glycero-3-phosphoethanolamine; EMA, European Medicines Agency; EPR, enhanced permeability retention; Eu.DOTA.DSA, europium(III) 2-{4,7-biscarboxymethyl-10-[(N,N-distearylamidomethyl-N′-amidomethyl]-1,4,7,10-tetraazacyclododec-1-yl}acetic acid; FDA, Food and Drug Administration; HIFU, high intensity focused ultrasound; LTSL, low temperature sensitive liposomes; MRI, magnetic resonance imaging; MSPC, 1-stearoyl-sn-glycero-3phosphocholine; NIR, near-infrared; PEG, polyethylene glycol; rf, radiofrequency; TSL, thermosensitive liposome; TTSL, traditional, temperature-sensitive liposomes; XL750.DSA, N′XenoLight750-N,N-distearylamidomethylamine

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DOI: 10.1021/mp5002679 Mol. Pharmaceutics XXXX, XXX, XXX−XXX