WO2012055020A1 - Thermosensitive liposomes - Google Patents

Abstract

In one aspect, the present invention provides a thermosensitive liposome comprising a lipid bilayer comprising 1,2-dipalmitoyl-sn-glycero-3-phosphatidylcholine (DPPC) and a compound of formula (I): C₁₇H₃₅(CH₂)_p(CO)_q(OCH₂CH₂)_nOH wherein p is an integer selected from 0 or 1; q is an integer selected from 0 or 1; p + q = 1; and n is an integer selected from about 10 to about 100.

Description

THERMOSENSITIVE LIPOSOMES

CROSS-REFERENCE TO RELATED APPLICATIONS

This application claims priority to U.S Patent Application Serial No. 61/407,250, filed October 27, 2010, the contents of which are hereby incorporated herein by reference in its entirety.

FIELD OF THE INVENTION

This invention relates to the field of liposomes and, more specifically, to thermosenstive liposomes capable of releasing an entrapped agent when heated.

BACKGROUND OF THE INVENTION Nanoparticle (NP)-based drug delivery systems including liposomes enable targeting of anticancer drugs to tumors, and their development and optimization have been a major focus in the field of drug delivery. Anticancer drugs encapsulated in liposomes have been demonstrated to improve the therapeutic window by enhancing the anti-tumor efficacy and reducing side effects [1]. Furthermore, the blood circulation time of liposomes can be prolonged by modification with polyethylene glycol (PEG), which acts to reduce the uptake of the delivery vehicle by the mononuclear phagocyte system (MPS) [2-3]. The prolonged circulation time of the liposomes leads to their increased accumulation in solid tumors by the enhanced permeability and retention (EPR) effect, wherein the liposomes enter the tumor via a leaky vasculature, and are not easily eliminated due the compromised lymphatic system [4]. PEGylated liposomal doxorubicin (DOX) (Doxil^®/Caelyx^®) has been approved clinically for Kaposi sarcoma, multiple myeloma and advanced ovarian cancer. However, while Doxil^® has minimized the acute cardiotoxicity associated with free DOX, it does not substantially increase the efficacy compared to the free DOX in the clinical setting. It is now widely understood that release of DOX from Doxil^® is slow (<5% in 24 h), leading to limited bioavailability (40-50%) [1, 5-7]. A similar problem is encountered with liposomal cisplatin (CDDP) (SPI-077), which demonstrated substantial tumor accumulation but displayed no antitumor activity in clinical trials [8]. Again, the release of the membrane impermeable CDDP was confirmed to be minimal from the liposomal formulation [9]. Moreover, the above described NPs have been found only to accumulate around the blood microvessels in the tumor with limited penetration and drug exposure [10], a potential cause of disease recurrence and resistance development. To overcome the release and penetration issues, attention in the drug delivery field has focused on designing NPs capable of releasing a drug when exposed to a specific triggering mechanism [1 1-13], which includes decreased pH in the tumor microenvironment [14], light [15], ultrasound [16-17], enzymatic action [18] or heat [19-21]. Among the trigger-sensitive NP formulations that have been developed, thermosensitive liposomes are the most advanced in commercial and clinical development, as focused induction of hyperthermia in deep tissue is clinically feasible and permissible [22]. Precise and localized heating can be achieved by application of technologies including radiofrequency ablation [23], focused microwave [24], high intensity focused ultrasound [25] and magnetic resonance guided focused ultrasound (MRgFUS) [26-27]. Studies have demonstrated that the use of hyperthermia in combination with the thermosensitive liposomal delivery resulted in increased intratumoral free drug concentrations compared to levels achieved with free drug or liposomal drug administered in the absence of heating, leading to significantly enhanced antitumor efficacy [21, 25, 28-59]. One of the most advanced thermal sensitive liposomal formulation is composed of DPPC/MSPC/DSPE-PEG (90/10/4, molar ratio), described as lysolipid-temperature-sensitive-liposomes (LTSL) [19, 21-22, 25, 28-29, 60-61] and is currently in Phase III clinical trials for liver cancer and Phase II for recurrent breast cancer on chest wall. When heated to 42°C, the LTSL released 100% DOX in 2-3 min [25, 60-63], leading to a ~15-fold increase in drug delivery to the heated tumor [21] and eradication of the s.c. inoculated human xenograft tumor in a mouse model [21, 28].

The LTSL-hyperthermia approach is differentiated from the conventional liposomal therapy by its non-dependence on the EPR effect. The thermosensitive liposomes have been administered during the hyperthermia treatment, with immediate release of the encapsulated drug to the heated tumor [64]. The liposomes have the effect of keeping the drug concentrated in the blood circulation by reducing the renal clearance, and as the liposomes circulate within a hyperthermic tumor, the drug is quickly released, generating a high drug concentration gradient, and driving diffusion from the blood into the tumor. In other words, the LTSL approach does not depend on the liposomes extravasating into the tumor through a leaky vasculature, a process which is not efficient, especially in hypovascular tumors [65]. Furthermore, the LTSL/hyperthermia strategy is particularly advantageous for the delivery of toxic drugs to localized and inoperable tumor, for which there are currently limited treatment options with disappointing therapeutic results [22]. There is a continued need for improved drug and therapeutic delivery systems.

SUMMARY OF THE INVENTION

According to a first aspect, there is provided a liposome comprising a lipid bilayer comprising l,2-dipalmitoyl-s«-glycero-3-phosphatidylcholine (DPPC) and a compound of formula: C i₇H₃₅(CH₂)_p(CO)_q(OCH₂CH₂)_nOH

(I) wherein

p is an integer selected from 0 or 1 ;

q is an integer selected from 0 or 1 ;

p + q = 1 ; and

n is an integer selected from about 10 to about 100.

Preferably, the lipid bilayer consists essentially of DPPC and Brij78, and further preferably consists of DPPC and Brij78.

According to a further aspect, there is provided a method of delivering an active agent to a target area in a patient, comprising administering to the patient a liposome described herein having an active agent entrapped in the interior space thereof; and heating the target area to at least 40°C for at least 3 min. In other embodiments, the target area is heated to at least 42°C for at least 3 min. According to a further aspect, there is provided a use of a liposome described herein for delivering an active agent to a target area in a patient, wherein the active agent is entrapped in the interior space of the liposome.

According to a further aspect, there is provided a use of a liposome described herein in the preparation of a medicament containing an active agent to be delivered to a target area in a patient, wherein the active agent is entrapped in the interior space of the liposome.

In still another aspect, there is provided a method for treating breast cancer, pancreatic cancer, or lung cancer in a patient in need thereof comprising administering to said patient a therapeutically effective amount of a liposome described herein, wherein an anticancer agent is entrapped in the interior space of the liposome. In another aspect, the anticancer agent is doxorubicin. In another aspect, the anticancer agent is gemcitabine.

In still another aspect, there is provided a method for treating multidrug resistant cancer in a patient in need thereof comprising administering to said patient a therapeutically effective amount of a liposome described herein, wherein an anticancer agent is entrapped in the interior space of the liposome. In another aspect, the anticancer agent is doxorubicin. In another aspect, the anticancer agent is gemcitabine.

In another aspect, the liposome further comprises a targeting moiety. In yet another aspect, the targeting moiety is an RGD targeting sequence. In still another aspect, the targeting moiety is covalently bound to the compound of formula I.

In yet another aspect, there is provided a liposome comprising a lipid bilayer comprising l,2-dipalmitoyl-s«-glycero-3-phosphatidylcholine (DPPC) and a compound selected from Brij 56, Brij 76, Brij 700, Myrj52, Myrj53, or Myrj59. In other aspects, the lipid bilayer consists essentially of DPPC and a compound selected from Brij 56, Brij 76, Brij 700, Myq^'52, Myrj53, or Myrj59. In other aspects, the lipid bilayer consists of DPPC and a compound selected from Brij 56, Brij 76, Brij 700, Myrj52, Myrj53, or Myrj59. BRIEF DESCRIPTION OF THE DRAWINGS

Embodiments of the invention may best be understood by referring to the following description and accompanying drawings. In the description and drawings, like numerals refer to like structures or processes. In the drawings: Figure 1 shows the chemical structures of selected Brij surfactants: (A) Polyoxyethylene (10) stearyl ether (Brij 76). (B) Polyoxyethylene (20) stearyl ether (Brij78). (C) Polyoxyethylene (20) oleyl ether (Brij98). (D) Polyoxyethylene (100) stearyl ether (Brij700).

Figure 2 shows the membrane permeability of the liposomes. (A) The absorbance decay of NBD in the LTSL formulation over time at 42°C. The NBD-labeled LTSL sample was put on a water bath at 42°C, and at selected time points, samples were taken and the absorbance was measured. (B) The membrane permeability rate constant for different liposomal formulations (For the Brij-formulations, DPPC/Brij-surfactant = 96/4, molar ratio). The rate constant was obtained by fitting the curve into the formula: Relative absorbance = mi*exp[-m₂*(t)] The m₂ is the membrane permeability rate constant. Data are mean ± S.D. (n=3).

Figure 3 shows the optimization of the drug loading conditions for the HaT formulation. DOX and the HaT-liposomes were mixed at various drug-to-lipid ratios (0.025, 0.05, 0.01, w/w), and the mixtures were incubated for 0-150 min at 37°C. Encapsulation efficiency of DOX into the liposomes was determined as described in the Materials and Methods (M&M) section. Data are mean ± S.D. (n=3).

Figure 4 shows DOX release from different liposomal formulations at different temperatures. The DOX containing liposomal formulations (LTSL: DPPC/MSPC/DSPE-PEG = 86/10/4; Brij-formulations: DPPC/Brij-surfactant = 96/4, molar ratio) were incubated at different temperatures and the drug release was measured at different time points as described in M&M section. Data are mean ± S.D. (n=3). Figure. 5 shows GEM and CDDP release from the HaT formulation at different temperatures. The GEM (A) and CDDP (B) containing HaT formulation was incubated at different temperatures and the released drug was isolated using ultrafiltration and was measured by LC/MS (A) and the colorimetric assay (B) as described in M&M. Data are mean ± S.D. (n=3).

Figure. 6 shows release of DOX from HaT formulated with differing molar ratios of Brij 78. The DOX containing formulations were incubated at 37°C or 42°C and the drug release was determined at different time points using the method described in M&M. Data are mean ± S.D. (n=3). ***, <0.005, significant difference compared to the DPPC-liposomes (0 mol% Brij78).

Figure. 7 shows Ti map of the Gd³⁺ containing liposomal formulations after different treatments. The formulations were incubated at 37, 40 or 42°C for 3 min and cooled down on ice before transferred to a 96-well plate for MRJ (NTSL: non-thermosensitive liposomes; HSPC/cholesterol/DSPE-PEG = 3/1/1, weight ratio). Triton X was used to disrupt the liposomes for complete release of Gd .

Figure. 8 shows the release of DOX from the Brij78-liposomes prepared by the thin film hydration method or the post-insertion method. The DOX containing formulations were incubated at 37 or 42°C for 30 min and the drug release was measured using the method described in M&M. Data are mean ± S.D. (n=3). *, /?<0.05, ***, ^<0.005, significant difference between the two methods.

Figure. 9 shows the intracellular uptake of DOX released from different liposomal formulations after different pre-treatments. EMT-6 cells were incubated at 37°C for 4 h with the formulations that had been preheated at 37, 40 or 42°C for 3 min. Intracellular uptake was quantified (A) using the method described in M&M or imaged by fluorescence microscopy (B-D). For A, Data are mean ± S.D. (n=3). **, p<0. \, significant difference compared to the data obtained at 37°C. For B-D, cells were counterstained with DAPI for nucleus. Images are 630* magnification, scale bars = 20 μηι. Figure. 10 shows the cytotoxicity of DOX in different formulations pre-incubated at different temperatures. DOX in different formulations (30 μΜ DOX) were pre-heated for 3 min at 37, 40 or 42°C in a water bath, and these solutions were applied to the EMT-6 cells for 4 h. The cell viability was determined 48 h later by the MTS assay. Data are mean ± S.D. (n=3).

Figure. 11 shows the pharmacokinetics of DOX in different formulations. DOX in different formulations was i.v. administrated (3 mg DOX/kg) into the tumor-free BALB/c mice. After the administration, the blood was collected at 15 min, 30 min and 1 h. Plasma level of DOX was determined as described in M&M. Data are mean ± S.D. (n=3). ***, ^><0.005, significant difference compared to free DOX.

Figure 11.1 shows drug release profiles and pharmacokinetics of different liposomal formulations. (A-C) Temperature dependent release of the HaT and LTSL formulations prepared with the pH gradient (labeled with Cit) or Cu²⁺ gradient method (labeled with Cu). HaT prepared with Cu²⁺ gradient [HaT(Cu)] was post-inserted with 0-32 mol% Brij78 by incubating the preformed particles with Brij78 at 37°C for 1 h. Un-incorporated Brij78 was then removed by gel filtration. The release assay was performed in HEPES buffered saline (HBS, pH 7.4). (D) Drug release from different liposomal formulations in HBS containing 30% fetal bovine serum at 37°C. (E) Pharmacokinetics of different liposomal formulations in healthy BALB/c mice. Figure. 12 shows the body weight change after the injection of DOX in different formulations. DOX in different formulations were i.v. administrated into the tumor-free BALB/c mice at 3 mg (A) or 10 mg DOX/kg (B). The body weight was monitored for the first week. Data are mean ± S.D. (n=3).

Figure. 13 shows an analysis of hemolytic activity. Murine red blood cells and various concentrations of liposomes were mixed and incubated at 37°C for 60 min. Triton X was used as a positive control. Hemolysis activity was measured as the increase of absorbance at 485 nm. Data are mean ± S.D. (n=3). Figure. 14 shows the biodistribution of DOX at 1 h after the i.v. injection of DOX in different formulations in combination with local hyperthermia (42°C) to the tumor in the left lower leg. The EMT-6 tumor-bearing mice received local hyperthermia (42°C) to the tumor in the left leg for 10 min before the i.v. injection of DOX in different formulations at the dose of 3 mg DOX/kg. The heating was maintained for another h before the sacrifice of the animals for measurement of drug concentrations in different tissues. The tumor in the right leg served as the unheated control. Data are mean ± S.D. (n=5-6). *, 7<0.05, **,p< .0\, ***,/?<0.005, significant difference compared with free DOX solution. #, p<0.05 ,/?<0.01, ^m, pO.Q05, significant difference compared with LTSL.

Figure 15 shows histological analysis of tumor growth 5 days after receiving different treatments. Local hyperthermia (42°C) was given to the tumor in the left leg of the EMT-6 tumor-bearing mice for 10 min before the i.v. injection of DOX in different formulations at the dose of 3 mg DOX/kg. The heating was maintained for another h and the tumor in the right leg served as the unheated control. Five days after the treatment of free DOX (A), LTSL-DOX (B) and HaT-DOX (C), the mice were sacrificed and the tumors on both legs were collected for H&E histological examination. The upper panel is 20x magnification, and the lower panel is 200x magnification. Alternatively, the tumor size was measured three times weekly after the treatment by caliper. Data are mean ± S.D. (n=5-6). ***, <0.005, statistically significant, heated vs unheated,†, ><0.05,††,/?<0.01, statistically significant compared with free DOX (heated), #, p<0.05, statistically significant compared with LTSL (heated).

Figure. 16 shows cytotoxicity of DOX with different formulations on A2780 cells and A2780-ADR cells. Various concentrations of DOX and liposomes containing solutions were applied to the A2780 cells (A) and A2780-ADR cells (B) for 24 h. The ratio of DOX and lipid was fixed (1 (μΜ)/15 (μΜ) corresponding to approximately 1/20 (w/w)). The cell viability was determined by the MTS assay in another 24 h after the drug treatment. Data are mean ± S.D. (n=3). Figure 17 shows the growth curve of individual tumors after different treatments. Local hyperthermia (42°C) was given to the tumor in the left leg of the EMT-6 tumor- bearing mice for 10 min before the i.v. injection of DOX in different formulations at the dose of 3 mg DOX/kg. The heating was maintained for another hour before the release of the animals. The mice only received a single treatment and the tumor in the right leg served as the unheated control. The tumor size was measured three times weekly after the treatment by caliper. Data are mean ± S.D. (n=4-5). *, <0.05, ***, <0.005, significant difference compared with the PBS only control. ^## , ><0.005, significant difference compared with Free DOX + hyperthermia.^†, ><0.05, significant difference compared with LTSL-DOX + hyperthermia.

Figure 18 shows the characteristics of various Brij and Myrj surfactants in DOX- loaded DPPC Liposomes, and the characteristics of DOX-loaded Brij 78 Liposomes with different phospholipids.

Figure 19 shows the MW distribution of the Brij 78 molecule, measured by ES- MS. Peak separations are 44.06 Da, corresponding to the CH₂CH₂0 repeating PEG units.

Figure 20 shows drug release for Brij formulations composed of

Ci₈H37(0CH₂CH₂)_n0H, wherein n = 2-100 (average numbers provided by

manufacturer). At n = 5 PEG repeat units (NMR estimation), thermosensitive properties are introduced. Figure 21 illustrates the synthesis scheme for a Brij molecule terminated with an cRGDPK peptide. This peptide sequence is intended to promote retention of the HaT molecule in tumor vasculature.

Figure 22 shows GEM release from the HaT and LTSL formulation at different temperatures in FBS-containing solutions. The HaT GEM (A) and LTSL GEM (B) formulations were incubated at different temperatures and the released drug was isolated using ultrafiltration and was measured by HPLC. Data are mean ± S.D. (n=3). The release data was fitted to the first order kinetics equation F = 1 - e-kt (where F = fraction of drug released, t = time, k = the drug release rate constant) and values for k were calculated. By this analysis (C) GEM HaT has an 8-fold higher release constant at 40-41°C.

Figure 23 shows the (A) pharmacokinetic (PK) profile of GEM (120 mg/kg), HaT GEM (20 mg GEM/kg) and LTSL GEM (20 mg/kg) in mice. In (B) is shown the biodistribution of GEM in plasma and tumor tissue.

Figure 24 shows the antitumor efficacy of GEM (120 mg/kg), HaT GEM (20 mg GEM/kg) and LTSL GEM (20 mg/kg) in mice bearing (A) PAN02 and (B) LL2 footpad tumors that have been heated at 42°C for 1 hour.

DETAILED DESCRIPTION OF THE PREFERRED EMBODIMENTS

In the following description, numerous specific details are set forth to provide a thorough understanding of the invention. However, it is understood that the invention may be practiced without these specific details. There is provided herein a simple and sensitive liposomal formulation for drug targeting and release under mild hyperthermia (40-42°C). The presently described liposomal formulation focuses on a minority component of Brij surfactants (preferably 1-8 mol%) combined with DPPC lipid. Brij molecules are single chain surfactants, commercially available with different PEG chain lengths and/or acyl chain structures (Fig. 1). A series of Brij containing liposomes was formulated and their sensitivity to mild hyperthermia conditions was examined. Brij compounds have been studied in colloidal formulations [66-67]. Moreover, Brij surfactants typically are nontoxic and common in pharmaceutical formulations [68], with safety being demonstrated in clinical trials [69]. In this study, a variety of in vitro analyses were performed to evaluate the temperature-dependent drug release. In vivo efficacy studies demonstrated that this novel formulation significantly improved drug delivery to the heated tumor and is more effective in regressing the solid tumor compared to free drug and the LTSL formulation. According to a first embodiment, there is provided a liposome comprising a lipid bilayer comprising l,2-dipalmitoyl-sn-glycero-3 -phosphatidylcholine (DPPC) and a compound of formula:

Ci₇H₃₅(CH₂)_p(CO)_q(OCH₂CH₂)_nOH

(I) wherein

p is an integer selected from 0 or 1 ;

q is an integer selected from 0 or 1 ;

p + q = 1 ; and

n is an integer selected from about 10 to about 100.

In some embodiments, n is an integer selected from 10, 15, 20, 25, 30, 35, 40, 45, 50, 55, 60, 65, 70, 75, 80, 85, 90, 95 or 100.

Preferably, the lipid bilayer consists essentially of DPPC and Brij78, and further preferably consists of DPPC and Brij78. As used herein, "liposomes" refer to artificially prepared vesicles made of lipid bilayer comprising a thin membrane made of two layers of lipid molecules.

In certain embodiments, the diameter of the liposome is from about 30 nm to about 250 nm.

In other embodiments, the liposome comprises at least one active agent entrapped in the interior space of the liposome, preferably a diagnostic agent or therapeutic agent. Preferably, the active agent is hydrophilic. More preferably, the active agent is hydrophilic and amphipathic.

The term "therapeutic agent" is art-recognized and refers to any chemical moiety that is a biologically, physiologically, or pharmacologically active substance. Examples of therapeutic agents, also referred to as "drugs", are described in well-known literature references such as the Merck Index, the Physicians Desk Reference, and The Pharmacological Basis of Therapeutics, and they include, without limitation, medicaments; vitamins; mineral supplements; substances used for the treatment, prevention, diagnosis, cure or mitigation of a disease or illness; substances which affect the structure or function of the body; or pro-drugs, which become biologically active or more active after they have been placed in a physiological environment. Various forms of a therapeutic agent may be used which are capable of being released from the subject composition into adjacent tissues or fluids upon administration to a subject.

In specific embodiments, the therapeutic agent is an anticancer agent such as: 5- imidodaunomycin, actinomycin, alemtuzumab, aminolevulinic acid, anastrozole, arninopterin, BBR3464, bevacizumab, bleomycin, busulfan, camptothecin, capecitabine, carboplatin, carinomycin, carmustine, cetuximab, Chlorambucil, chlormethine, cisplatin, cladribine, clofarabine, cyclophosphamide, cytarabine, dacarbazine, dasatinib, daunorubicin, denileukin diftitox, docetaxel, doxorubicin, epirubicin, erlotinib, etoposide, exemestane, fiuorouracil, floxuridine, fludarabine, fotemustine, fulvestrant, gefitinib, gemcitabine, gemtuzumab, genasense, goserelin acetate, idarubicin, ifosfamide, imatinib, irinotecan, lapatinib, letrozoleleuprolide acetate, lomustine, mechlorethamine, megestrol acetate, melphalan, mercaptopurine, methotrexate, methyl aminolevulinate, mitomycin, mitoxantrone, N30 acetyldaunomycin, nacetyladriamycin, nilotinib, oxaliplatin, paclitaxel, panitumumab, pemetrexed, pentostatin, porfimer sodium, procarbazine, raltitrexed, rituximab, rubidazone, sorafenib, streptozocin, sunitinib, tamoxiifen, TbioTEPA, temozolomide, teniposide, testolactone, thioguanine, topotecan, toremifene, tositumomab, trastuzumab, triptorelin pamoate, uramustine, valrubicin, vandetanib, velcade, verteporfin, vinblastine, vincristine, vindesine, and vinorelbine.

In other embodiments, the therapeutic agent may be selected from ellipticine, prednisone, methyl-prednisone and ibuprofen.

A "diagnostic" or "diagnostic agent" is any chemical moiety that may be used for diagnosis. For example, diagnostic agents include imaging agents, such as those containing radioisotopes such as indium or technetium; contrasting agents containing iodine or gadolinium; enzymes such as horse radish peroxidase, GFP, alkaline phosphatase, or -galactosidase; fluorescent substances such as europium derivatives; luminescent substances such as N-methylacrydium derivatives or the like.

In some embodiments, the active agent is loaded into the liposome via a pH gradient. In other embodiments, the molar ratio of DPPC to the compound of formula I is 99:1 to 92:8, and in various embodiments, is at least 99:1, at least 98:2, at least 96:4, about 96:4 or at least 92:8. Preferably, the active agent is mixed with the liposome suspension at an active agen lipid ratio of 1 :5 to 1 :40 w/w.

In still other embodiments, the active agent is loaded into the liposome via a copper ion gradient. In other embodiments, the molar ratio of DPPC to the compound of formula I is 84:16 to 68:32. In other embodiments, the active agent is mixed with the liposome suspension at an active agent: lipid ratio of 1 :5.

In some embodiments, the liposome described herein has a T_m of about 42°C.

In some embodiments, there is less than 20% leakage of the active agent at 37°C in 30 min, preferably less than 5% leakage of the active agent at 37°C. In some embodiments, there is substantially full release of the active agent in less than 10 min, preferably in less than 3 min, when the liposome is heated to 40°C.

In other embodiments, there is substantially full release of the active agent in less than 10 min, preferably in less than 3 min, when the liposome is heated to 42°C.

In some embodiments, the liposome described herein is for delivery of an active agent to a target area in a patient, wherein the active agent is entrapped in the interior space of the liposome.

According to a further embodiment, there is provided a method of delivering an active agent to a target area in a patient, comprising administering to the patient a liposome described herein having an active agent entrapped in the interior space thereof; and heating the target area to at least 40°C for at least 3 min. In other embodiments, the target area is heated to at least 42°C for at least 3 min. According to a further embodiment, there is provided a use of a liposome described herein for delivering an active agent to a target area in a patient, wherein the active agent is entrapped in the interior space of the liposome.

According to a further embodiment, there is provided a use of a liposome described herein in the preparation of a medicament containing an active agent to be delivered to a target area in a patient, wherein the active agent is entrapped in the interior space of the liposome.

In still another embodiment, there is provided a method for treating breast cancer, pancreatic cancer, or lung cancer in a patient in need thereof comprising administering to said patient a therapeutically effective amount of a liposome described herein, wherein an anticancer agent is entrapped in the interior space of the liposome. In another embodiment, the anticancer agent is doxorubicin. In another embodiment, the anticancer agent is gemcitabine.

In still another embodiment there is provided a method for treating multidrug resistant cancer in a patient in need thereof comprising administering to said patient a therapeutically effective amount of a liposome described herein, wherein an anticancer agent is entrapped in the interior space of the liposome. In another embodiment, the anticancer agent is doxorubicin. In another embodiment, the anticancer agent is gemcitabine. In another embodiment, the liposome further comprises a targeting moiety. In yet another embodiment, the targeting moiety is an RGD targeting sequence. In still another embodiment, the targeting moiety is covalently bound to the compound of formula I.

In yet another embodiment, there is provided a liposome comprising a lipid bilayer comprising 1 ,2-dipalmitoyl-s«-glycero-3 -phosphatidylcholine (DPPC) and a compound selected from Brij 56, Brij 76, Brij 700, Myrj52, Myrj53, or Myrj59. In other embodiments, the lipid bilayer consists essentially of DPPC and a compound selected from Brij 56, Brij 76, Brij 700, Myrj52, Myrj53, or Myrj59. In other embodiments, the lipid bilayer consists of DPPC and a compound selected from Brij 56, Brij 76, Brij 700, Myrj52, Myrj53, or Myrj59.

The following examples are illustrative of various aspects of the invention, and do not limit the broad aspects of the invention as disclosed herein. EXAMPLES

MATERIALS AND METHODS

Materials

1 ,2-dipalmitoyl-5«-glycero-3-phosphatidylcholine (DPPC), l-stearoyl-2-hydroxy-s«- glycero-3 -phosphatidylcholine (MSPC), 1 ,2-distearoyl-sn-glycero-3- phosphatidylethanol-amine-N-[methoxy (polyethyleneglycol)-2000] (DSPE-PEG₂₀oo), 1 ,2-dihexadecanoyl-5n-glycero-3-phospho-(l'-rac-glycerol) (DPPG), 1 ,2- ditetradecanoyl-5«-glycero-3 -phosphoethanolamine (DMPE), L-a-phosphatidylcholine (HSPC), and 1 ,2,-dipalmitoyl-5«-glycero-3-phosphoethanolamine-N-(7-nitro-2-l ,3- benzoxadiazol-4-yl (NBD-PE) were purchased from Avanti Polar Lipids (Alabaster, AL). Polyoxyethylene stearyl ethers: Brij 30, 35 (L23), 52, 56 (CIO), 58, 72, 76 (CIO), 78 (S20), 93, 97 (O10), 98, and 700 (S100) were purchased from Sigma Aldrich (Oakville ON). Polyoxyethylene stearates: Myrj 45 and 59 were purchased from Lab Express (Fairfield NJ), and Myrj 52 and 53 were purchased from Sigma Aldrich (Oakville ON). Those skilled in the art will understand that equivalent compounds may be sold under different names (for instance, Croda, a supplier of Brij 78, has renamed this molecule Brij S20). DOX was purchased from Tocris Bioscience (Ellisville, MO). Gemcitabine (GEM) was purchased from BetaPharma (Branford, CT). Peptides were purchased from Peptides International (Louisville, KY). All other reagents were purchased from Sigma Aldrich (Oakville, ON) and were of analytical grade. Preparation of liposomes for the membrane permeability study

The liposomes (formulations listed in Table 1) with 1 mol% NBD-PE were prepared by the thin film hydration method, followed by membrane extrusion for size control as described previously [61]. Briefly, the lipids were dissolved in isopropanol (IP A) and dried under a stream of nitrogen gas, and the resulting thin lipid film was placed under high vacuum for at least 2 h to remove residual organic solvent. The lipid film was hydrated with a solution of 100 mM NaCl with 709 mM sucrose to form multilamellar vesicles (MLVs). MLVs were sonicated for 10 min and subsequently extruded 10 times through stacked polycarbonate filters of 0.1 μπι pore size at 65°C to adjust the size. The liposomes were then cooled to room temperature. The particle size of the liposomes was determined by dynamic light scattering with a particle analyzer (Zetasizer Nano-ZS, Malvern Instruments Ltd, Malvern, UK). The liposomal phospholipid concentration was quantified by the Fiske and Subbarow phosphate assay [70]. The diameter of different liposomal formulations is summarized in Table 1.

Table 1. Physicochemical properties of the liposomes.

DOX-loaded liposomes prepared by pH gradient method

The DOX-loaded liposomes were prepared using the method described previously with some minor modifications [61]. The thin lipid films (formulations listed in Table 1) were hydrated with 300 mM citric acid. After the sonication and membrane extrusion (at 65°C) to control the size, the liposomes were cooled to room temperature, and the exterior buffer of the liposomes was replaced by HBS (25 mM HEPES Buffered Saline, pH 7.4) via dialysis using a dialysis cassette (Slide-A-Lyzer 10 kDa MWCO, Pierce Biotechnology, Rockford, IL) for 3 h against three exchanges of 500X volume HBS. The liposome suspension and DOX were mixed at 1 : 10-1 :40 (w/w, drug/lipid), and the mixture was incubated at 37°C for various time points. After incubation, un- encapsulated DOX was removed by gel filtration on a Sepharose CL-4B column (Sigma- Aldrich, St Louis, MO) equilibrated with HBS. The eluted liposome fraction was analyzed for lipid and drug content. The liposomal phospholipid concentration was quantified by the Fiske and Subbarow phosphate assay [70]. The DOX concentration was determined by measuring the fluorescence after the disruption of liposomes with Triton X-100 using a Chameleon multilabel plate reader (Hidex Personal Life Science, Hidex Oy, Finland) (Ex 485 nm/Em 590 nm) and compared with a standard curve. The encapsulation efficiency was calculated as [DOX/lipid (after gel filtration)]/[DOX/lipid (before gel filtration)] x 100%. The diameter and encapsulation efficiency of different liposome formulations are summarized in Table 1. The Brij78-liposome (DPPC/Brij78 96:4 mol/mol), which was determined to be an optimal formulation, is referred to as the Hyperthermia-activated cytoToxic (HaT) formulation throughout this study.

Unless otherwise indicated, the DOX-loaded liposomes referred to herein were prepared by this method. Gemcitabine (GEM)-loaded HaT liposomes

A lipid film (DPPC/Brij 78 =96/4, molar ratio) was hydrated with 0.9% NaCl solution containing GEM (0.1 mg GEM / mg DPPC) at 65°C. The liposomes were controlled for size by membrane extrusion (100 nm) at 65°C, the solution was cooled to room temperature, and dialyzed against 500X volume of 0.9% NaCl with three exchanges over 3 h to remove the remaining un-encapsulated drug. The resulting HaT-GEM was diluted in saline with TritonlOO to disrupt the liposomes, and this solution was analyzed by a Waters HPLC system (acetonitrile/water gradient (0.1% formic acid), 5-95% acetonitrile over 6 min). The encapsulation efficiency of GEM by the passive loading procedure was 4-20%, and the particle size was measured to be 95-120 nm with a PDI < 0.1 (n=3).

Gemcitabine (GEM)-loaded LTSL liposomes

A lipid film (DPPC/MSPC/DSPE-PEG₂₀₀₀ = 86/10/4, molar ratio) was hydrated with 0.9% NaCl solution containing GEM (0.15 mg GEM / mg DPPC) at 65°C. The liposomes were controlled for size by membrane extrusion (100 nm) at 65°C, the solution was cooled to room temperature, and dialyzed against 500X volume of 0.9% NaCl with three exchanges over 3 h to remove the remaining un-encapsulated drug. The resulting HaT-LTSL was diluted in saline with TritonlOO to disrupt the liposomes, and this solution was analyzed by a Waters HPLC system (acetonitrile/water gradient (0.1% formic acid), 5-95% acetonitrile over 6 min). The encapsulation efficiency of GEM by the passive loading procedure was 4-20%, and the particle size was measured to be 95-120 nm with a PDI < 0.1 (n=3).

CDDP-loaded liposomes

A lipid film (DPPC/Brij 78 =96/4, molar ratio) was hydrated with 0.9% NaCl solution containing cisplatin (0.15 mg CDDP / mg DPPC) at 90°C. The liposomes were controlled for size by membrane extrusion (100 nm) at 90°C, the solution was cooled to room temperature, and centrifuged at 2,000 rpm for 5 min to remove un-encapsulated and precipitated CDDP. The resulting solution was dialyzed against a 500X volume of 0.9% NaCl with three exchanges over 3 h to remove the remaining un-encapsulated drug. The resulting HaT-CDDP was diluted in IPA to disrupt the liposomes, and this solution was analyzed by the colorimetric method for CDDP content [71]. The encapsulation efficiency of cisplatin by the passive loading procedure was 3.8%, and the particle size was measured to be 100- 120 nm with a PDI < 0.1. Differential scanning calorimetry (DSC)

DSC measurements were performed as described previously [72]. The liposomes prepared in 300 mM citric acid were frozen on dry ice and lyophilized for overnight with a lyophilizer (Freezone 4.5 freeze drier, Labconco, Kansas city, MO). The lyophilized liposomes (5-10 mg) were transferred to an aluminum pan and analysis was performed using a Q100 differential scanning calorimeter (TA Instruments, New Castle, DE). The samples were analyzed at a scan rate of 10°C/min between 20°C and 70°C. The data were analyzed using TA Universal software (TA Instruments, New Castle, DE). The melting point (T_m) was defined as the offset of the heat flow peak. The T_m of different liposomal formulations is described in Table 1, and ranged between 40-42°C.

Measurement of liposomal membrane permeability

Measurement of liposomal membrane permeability was performed according to a reported method [61]. Briefly, 200 μΐ of the NBD-labeled liposomes was gently mixed with 5 μΐ of dithionite solution (1M sodium dithionite with 1M TRIZMA buffer) and the mixture was kept at room temperature for 5 min. The sample was incubated at different temperatures (30, 37, 40 or 42°C) for various time points (1, 3, 5, 10, 15, 20, 25 or 30 min), and was immediately placed on ice, and subsequently transferred to a 96-well plate for the measurement of absorbance at 485 nm. The absorbance decay of the sample was plotted against the time and the curve was fitted to the equation: Relative absorbance = mi*exp[-m₂*(t)], where m₂ is the membrane permeability rate constant. The curve fitting was performed using GraphPad Prism (GraphPad Software Inc., San Diego, CA) In vitro release of DOX, GEM, and CDDP from liposomes

Measurement of DOX released from the liposomes was demonstrated as described previously [61]. The liposomes (1 μg DOX/ml in 200 μΐ HBS) were incubated at different temperatures (30, 37, 40 or 42°C) for various time points (1, 3, 5, 10, 15, 20, 25 or 30 min), and were immediately put on ice and transferred into a 96-well plate. The release of DOX was determined using a Chameleon plate reader by measuring the fluorescence (Ex 485 nm/Em 590 nm). The percentage of the released DOX was calculated as (Ιχ-Ιο)/(Ιιοο-Ιο) ^x 100%, in which I_T is the fluorescence at time point t, I₀ is the fluorescence at the start of the incubation time, l_m is the fluorescence after the addition of 10 μΐ of 0.5% Triton X-100.

The HaT-GEM formulation was diluted with 0.9% NaCl solution to adjust the concentration to 20 μg GEM/ml and was incubated at 20, 37, 40 or 42°C, for 1, 2, 5, 10, 20 or 30 min, in triplicate. At selected time points, the sample was placed on ice for 5 min, transferred to an Amicon Ultra centrifugal filter unit (MWCO = 10K), and spun for 5 min at 14,000 rpm. An aliquot of the filtrate (released drug) was then analyzed for GEM concentration by LC/MS.

The HaT-CDDP formulation was diluted with 0.9% NaCl solution to adjust the concentration to 50 μg CDDP/ml and was incubated at 20, 37, 40 or 42°C, for 1, 2, 5, 10, 20 or 30 min, in triplicate. At selected time points, the sample was placed on ice for 5 min, was transferred to an Amicon Ultra centrifugal filter unit (MWCO = 10K), and spun for 5 min at 14,000 rpm. An aliquot of the filtrate (released drug) was then analyzed for CDDP concentration by the colorimetric assay [71]. Stability of DOX-loaded HaT liposome

The stability of the DOX-loaded liposomes was measured as described previously [70]. The liposomes were stored in the dark at 4°C and at selected time points (1 day, 1 week, 2 weeks and 1 month), liposome samples were diluted with HBS, the diameter was measured by the Zetasizer Nano-ZS and the leakage of the drug was determined using the method for DOX release described earlier. Detection of Gd-DTPA release from liposomes by magnetic resonance imaging (MRP

The detection of Gd-DTPA release was demonstrated as described previously [62]. Briefly, the liposomes were hydrated with 150 mM gadolinium-diethylenetriamine penta-acetic acid (Gd-DTPA), were membrane extruded at 65°C for size control, cooled down to room temperature, and dialyzed to remove free Gd-DTPA using a 500X volume of HBS. The liposomes (200 μΐ, 0.8-1 mM Gd-DTPA) were incubated at different temperatures (37, 40 or 42°C) for 3 min, followed by chilling on ice, and were transferred into a 96-well plate. MRI monitoring of Gd-DTPA release was performed using a 7T micro-MRI (BioSpec 70/30 USR, Bruker Biospin, Ettlingen, Germany) located at the STTARR facility (Radiation Medicine Program, University of Toronto, Ontario, Canada). Tl maps were obtained using a saturation-recovery technique, involving acquisition of Tl -weighted RARE images at variable repetition times (TE = 8.7 ms; TR = 25, 50, 100, 150, 250, 500, 750, 1000, 1500, 2500, 5000, 7500 ms; 400 x 500 μηι in-plane resolution over 90x80 mm field-of-view; 2 mm slice thickness; RARE factor 2; 19 minute data acquisition). The mean and standard deviation of Tl for voxels within each well was calculated via histogram analysis within manually-traced ROIs (Mipav software, National Institutes of Health, Bethesda, MD, USA).

Preparation of DOX containing DPPC/Brij 78-liposomes using the post-insertion method The post-insertion method to incorporate Brij78 into the DPPC-liposomes was demonstrated as described previously [73]. Briefly, the DPPC-liposomes hydrated with 300 mM citric acid were gently mixed with an aliquot of the Brij78 solution (1, 2, 4 or 8 mol%, prepared in 300 mM citric acid), and incubated at 37°C for 1 h. The complete insertion of the Brij78 into the liposomes was confirmed by the disappearance of the micelle peak of Brij78 (10 nm) determined by the Zetasizer Nano-ZS. The obtained liposomes were dialyzed against HBS, and DOX was loaded into the liposomes as described previously. DOX release from the liposomes was measured as described previously. Cell culture

The mouse mammary carcinoma cell line EMT-6 was a generous gift from Dr. David Stojdl at the CHEO Research Institute and Dr. Douglas Mahoney at the University of Ottawa. EMT-6 cells were maintained in DMEM supplemented with 10% heat- inactivated fetal bovine serum, penicillin (100 U/ml) and streptomycin (100 μg/ml) at 37°C with 5% C0₂. The human ovarian carcinoma cell line, A2780 and adriamycin- resistant cell line (A2780-ADR) were obtained from Dr. Jeremy Squire at Queen's University, and were maintained in RPMI 1640 supplemented with 10% heat- inactivated fetal bovine serum, penicillin (100 U/ml) and streptomycin (100 μg/ml) at 37°C with 5% C0₂. The A2780-ADR cells were maintained with 0.1 μΜ DOX. PAN02 murine pancreatic and LL/2 murine lung cancer cells were maintained in DMEM supplemented with 10% heat-inactivated fetal bovine serum, penicillin (100 U/ml) and streptomycin (100 μ^πιΐ) at 37°C with 5% C0₂. Intracellular uptake of DOX released from the liposomes

The experiment was performed according to the method reported with minor modifications [74]. EMT-6 cells were seeded at a density of 2.5 x 10⁴ cells per well in 500 μΐ culture medium in a 24-well plate. Twenty-four h later, the medium was replaced with the liposome suspensions that had been diluted with the culture medium to a concentration of 30 μΜ DOX and heated at 37, 40, or 42°C for 3 min. The cells were incubated for 4 h at 37°C, after which the cells were gently washed two times with chilled PBS, and lysed by the addition of 0.5 ml of 0.3 % Triton-X/PBS solution with agitation on a rotating platform for 15 min at room temperature. Acidified IPA (75 mM HC1, 10% water/90% IPA) (1.5 ml/well) was added to the lysate and the mixture was incubated at 4°C in the dark for overnight. The cell lysate was collected and centrifuged for 3 min at 12,000 x g, and the supernatant was analyzed for the fluorescence intensity using a plate reader (Ex 485 nm/Em 590 nm). The protein content of the lysate was measured with the protein assay kit (Bio-Rad Laboratories, Hercules, CA), based on the Bradford method [75]. The data of intracellular uptake of DOX are expressed as fiuorescence/mg protein. Intracellular uptake of DOX was also observed by fluorescence microscopy. EMT-6 cells were seeded at a density of 2.5 * 10⁴ cells/coverslip (12 mm* 12 mm round), in 500 μΐ medium in a 24-well plate. After the treatment described above, the cells were fixed in 1% formalin in PBS for 15 min and then stained with 4', 6-diamidino-2- phenylindole (DAPI, 0.5 μ^πιΐ in PBS) for 10 min. After rinsing the slide with PBS, the coverslip was mounted on a glass slide. The cells were imaged by a fluorescent microscope (Axio Observer Zl, Carl Zeiss, Gottingen, Germany) with the Axiovision software (Carl Zeiss).

Cytotoxicity assay The cytotoxicity assay was demonstrated as described previously with minor modifications [62]. The EMT-6 cells were seeded at a density of 5 x 10 cells per well in 100 μΐ medium in a 96-well plate. At 24 h after seeding, the medium was replaced with various concentrations of preheated DOX formulations (37, 40 or 42°C for 3 min). Four h after the incubation at 37°C, the medium was removed and then the cells were gently washed with PBS twice. Subsequently, cells were incubated with 100 μΐ of fresh medium for 48 h. The cytotoxicity was determined by the MTS assay (Promega Celltiter 96^® AQueous Non-Radioactive Cell Proliferation Assay, Promega, Madison, WI), following the manufacturers' protocol. The IC50 was determined by nonlinear regression analysis using GraphPad Prism. Animals

The female BALB/c mice (aged 6 weeks, 18-20 g) were purchased from The Jackson Laboratory (Bar Harbor, ME). All experimental protocols in this study were approved by the Animal Care Committee of the University Health Network (Toronto, Ontario, Canada) in accordance with the policies established in the Guide to the Care and Use of Experimental Animals prepared by the Canadian Council of Animal Care. Pharmacokinetic study for HaT-DOX

LTSL-DOX, HaT-DOX or free DOX (3 mg DOX/kg) were i.v. injected into the tumor- free mice via the tail vein. At various time points (10 min, 30 min and 1 h) after the injection, the mice were placed under deep analgesia by administration of isoflurane, and the blood (100 μΐ) was collected by cardiac puncture. The blood sample was quickly transferred into an EDTA-coated collection tube (MiniCollect, Greiner Bio- One, The Netherlands). The plasma was isolated by centrifuging the blood samples at 4°C for 15 min at 2,500 rpm. The plasma level of DOX was measured by the method reported earlier [74]. Briefly, 10 μΐ of plasma was diluted with 990 μΐ acidified IPA and the mixture was incubated at 4°C in the dark for overnight. The sample was then centrifuged for 10 min at 12,000 x g and the supernatant was loaded onto a 96-well plate for fluorescence determination (Ex 485 nm/Em 590 nm) and the DOX concentration was calculated from a standard curve. The AUC was calculated using poly-exponential curve fitting and the least-squares parameter estimation program SAAMII (Micromath, UT).

Alternative method of preparing DOX-loaded liposomes - copper ion gradient method

DOX-loaded liposomes were prepared using a method alternate to the pH gradient method described above. Thin lipid films (DPPC/Brij 78 = 96/4) were hydrated with 100 mM CuS0₄ (pH 3.5), and were sized to ~ 100 nm by membrane extrusion. The external buffer of the liposomes was exchanged by passing the solution through a Sepharose CL-4B column equilibrated with a buffer composed of 300 mM sucrose, 20 mM HEPES, and 15 mM EDTA, with a pH of 7.4. The liposome suspension and DOX were mixed at 1 :5 (w/w, drug/lipid), and the mixture was incubated at 37°C for 80 minutes. The preformed liposomes were gently mixed with an aliquot of Brij 78 solution containing 0-32 mol% Brij 78 (compared to total lipid content), and incubated at 37°C for 1 hour. Un-encapsulated DOX and un-incorporated Brij 78 were removed by filtration through a Sepharose CL-4B column conditioned with HBS. The eluted liposome fraction was analyzed for lipid and drug content as described earlier. The drug release and PK profiles were analyzed as described earlier. Preliminary toxicity study for HaT-DOX

LTSL-DOX, HaT-DOX or free DOX (3 mg or 10 mg DOX/kg) were i.v. injected into the tumor-free mice via tail vein. The body weight was measured two or three times weekly after the drug administration. Hemolysis assay

The hemolysis assay was carried out as described previously [76]. Briefly, 10 μΐ of the murine red blood cells were diluted with 990 μΐ of HBS. The solution was mixed with various concentrations of liposomes, and was incubated at 37°C for 1 h. A Triton X- 100 treated red blood cell sample was included as a positive control. After the incubation, the solutions were centrifuged at 4°C for 10 min at 3,000 rpm and the supernatants were loaded onto a 96-well plate for the measurement of absorbance at 480 nm using a Chameleon plate reader.

Biodistribution study for HaT-DOX

The EMT-6 cells (1 x 10⁶ cells in 50 μΐ medium) were s.c. implanted into both lower legs of the mice. Seven days post-tumor inoculation (tumor mass was approximately 0.2-0.3 g), the mice were anesthetized by administration of isoflurane, and the left lower leg was taped onto a thermostatically controlled heating pad (FHC, Bowdoinham, ME) with the tumor in direct contact with the 42 ± 0.5°C surface. The tumor on the right leg was used as the unheated control. The tumor was heated for 10 min to equilibrate the temperature before the i.v. injection of different drug formulations (3 mg DOX/kg). The hyperthermia was maintained for 1 h after the injection. After the treatment, the blood (100 μΐ) was collected, the mice were sacrificed, and the heart, kidney, liver, lung, spleen and tumors were immediately excised. The tissue samples were washed with PBS and weighed after removing excess fluid. The DOX content in the tissues was determined using the method described previously [7]. Briefly, 0.1-0.3 g of the tissue samples were suspended in 1.5 mL nuclear lysis buffer (10 mM HEPES, 1 mM MgS0₄, 1 mM CaCl₂, pH 7.4) and homogenization was performed for 2 x 30 s at 6600 rpm with a tissue homogenizer (Precellys 24, Bertin Technologies, Cartland, CA). An aliquot of the homogenate (100 μΐ) was transferred into a 1.5 ml microtube, and 50 μΐ of 10% (v/v) Triton X-100, 100 μΐ of water, and 750 μΐ of acidified IPA were added and the mixture was stored for overnight at -20°C. The mixture was centrifuged for 10 min at 12,000 x g and the supernatant was loaded onto a 96-well plate (Ex 485 nm/Em 590 nm) for DOX determination. The data were compared with a standard curve made from spiking known amounts of DOX into the tissue homogenates from the untreated mice to get the absolute quantification of DOX in the tissue.

Antitumor efficacy study for HaT-DOX The EMT-6 cells (2 ^χ 10⁵ cells/50 μΐ medium) were s.c. inoculated into both lower legs of the mice. After 7 days, the tumor became palpable and the treatments described in the biodistribution study were initiated. The tumor size was measured using a caliper and the body weight was also monitored. Alternatively, 5 days after the treatment, 2-3 mice from each group were sacrificed and the tumor was collected for tissue section, H&E staining and histological analysis.

In vitro cytotoxicity study on the adriamycin drug resistant (ADR) cells

The cytotoxicity assay was demonstrated as described previously with minor modifications [77]. The A2780 cells and A2780-ADR cells were seeded at a density of 5 x 10³ cells per well in 100 μΐ medium in 96-well plates. Twenty-four h after seeding, the medium was replaced with solutions containing various concentrations of DOX and liposome formulations. The ratio of DOX and lipid content of each solution was fixed at 1/15 molar ratio, corresponding to an approximately 1/20 weight ratio. After incubation for 24 h at 37°C, the medium was removed from the wells, the cells were gently washed twice with PBS, and incubated with 100 μΐ of fresh medium for 24 h. The cytotoxicity was then determined by the MTS assay. Pharmacokinetic and Biodistribution study for HaT-GEM

LTSL-GEM (20 mg GEM/kg), HaT-GEM (20 mg GEM/kg) or free GEM (120 mg GEM/kg) were i.v. injected into EMT-6 s.c. tumor-bearing mice via the tail vein. At various time points (10 min, 30 min and 1, 2 and 4 h) after the injection, the mice were placed under deep analgesia by administration of isoflurane, and the blood (100 μΐ) was collected by cardiac puncture. The blood sample was quickly transferred into an EDTA-coated collection tube (MiniCollect, Greiner Bio-One, The Netherlands), and mixed with tetrahydrouridine (THU). The plasma was isolated by centrifuging the blood samples at 4°C for 15 min at 2,500 rpm. The plasma level of GEM was measured by HPLC. The AUC was calculated using poly-exponential curve fitting and the least- squares parameter estimation program SAAMII (Micromath, UT). One hour after dose administration the mice were sacrificed and the GEM uptake in the heated and unheated tumors, and the remaining dose in the plasma, were determined by LC/MS.

Efficacy study of HaT-GEM in Mice Bearing Footpad Tumors PAN02 or LL2 cells (2 10⁵ cells/50 μΐ medium) were s.c. inoculated into the right feet of mice. Seven days post-tumor inoculation (tumor mass was approximately 0.2- 0.3 g), the mice were anesthetized by administration of isoflurane, the mice were immobilized on a vertical surface, and their right tumor bearing footpad was immersed in a 42°C water bath. The tumor was heated for 10 min to equilibrate the temperature before the i.v. injection of different drug formulations (20 mg GEM/kg for the HaT and LTSL formulations, 120 mg GEM/kg for free GEM). The hyperthermia was maintained for 1 h after the injection. The tumor size at selected timepoints was measured using a caliper.

Statistical analysis All data are expressed as the mean ± SD. Statistical analysis was performed with the two-tailed unpaired t-test for 2-group comparison and/or one way ANOVA, followed by Tukey's multiple comparison test by using GraphPad Prism (for 3 or more groups). A difference with »<0.05 was considered to be statistically significant. Example 1: Membrane permeability of the liposomal formulations

In this study, the liposomal membrane permeability at different temperatures was determined. The dithionate ion, S₂0₄ ^2", is a reaction intermediate in the nitro reduction of NBD to its corresponding amine, resulting in the irreversible quenching of NBD fluorescence and decay of the absorbance peak of NBD at 465 nm [78]. Fig. 2A depicts the absorbance decay in the LTSL formulation (DPPDC/MSPC/DSPE-PEG) at 42°C. With the addition of the dithionite solution, about half of the NBD-fluorescence was quenched, an effect which indicates that about half of the NBD molecules in the liposomes were decorating the outer surface of the liposomes (Fig. 2A, time zero). The absorbance decay observed after the incubation of the LTSL formulation at 42°C indicates the permeation of dithionite ion to the interior space of the liposomes with quenching of the interior NBD-fluorescence molecules. To quantify the membrane permeability, we obtained the membrane permeability rate constant by fitting the curve of absorbance decay to an exponential equation as described in the Materials and Methods section.

We obtained the membrane permeability rate constant for 4 types of Brij -liposomes (DPPC/Brij -surfactant = 96/4), with the LTSL formulation as a positive control and the DPPC -liposome as a negative control. As shown in Fig. 2B, the Brij-liposomes (Brij78, Brij76 and Brij700) showed increased membrane permeability compared to the DPPC-liposomes at 40-42°C. Notably, the HaT formulation (Brij78-liposomes) showed further increased membrane permeability upon mild hyperthermia compared to the LTSL formulation (a 3.2-fold increase at 40°C and a 3.4-fold increase at 42°C). The Brij 76 formulation also exhibited increased thermal sensitivity compared to the LTSL at 42°C. The Brij98-liposomes, however, displayed no temperature dependant increase in membrane permeability.

Example 2: Optimization of the drug loading

into the Brij78-liposomes

We selected the Brij78-liposomes (DPPC/Brij 78 = 96/4) as the formulation to optimize for drug loading conditions, as it demonstrated the best thermal sensitivity among all formulations screened (Fig. 2). The LTSL formulation demonstrated 100% encapsulation efficiency at 60 min of incubation at 37°C, in contrast to the 90 min that was required for the HaT formulation. When the drug-to-lipid ratio of HaT was reduced from 0.05 to 0.025, the HaT formulation demonstrated 100% encapsulation efficiency at 60 min. When the drug-to-lipid ratio of HaT was 0.1 , only 40-50% drug loading was achieved. The optimal drug loading condition for the Brij78-liposomes therefore appeared to be at 37°C for 90-120 min with a drug to lipid ratio of 0.05 (Fig. 3), which is comparable with the LTSL formulation. Increasing the incubation time to 150 min decreased the drug loading to 70%. Accordingly, DOX was then loaded to the formulations using the optimal condition 90-120 min condition.

Example 3: Temperature dependent drug release from different liposomal formulations

A series of liposomes loaded with DOX were incubated at 30, 37, 40 or 42°C, and the drug release over a period of 30 min was measured. As depicted in Fig. 4 (A-E), the DOX release profile depended on the composition of the liposomes. The LTSL formulation released 100% DOX within 3 min at 42°C, but at 40°C only reached full release after 10 min incubation (Fig. 4A). The DPPC-only liposomes demonstrated a temperature dependent DOX release profile (Fig. 4B), but this formulation did not release 100% by the end of the experiment (72.5%, 30 min at 42°C), and furthermore, DOX leakage (31.9%, 30 min) was observed at 37°C. The Brij76-liposomes exhibited a rapid release profile (Fig. 4C) at 42°C, with full release within 3 min. However, the Brij76-liposomes exhibited a relatively slow release kinetics at 40°C (100% release in 15 min), and the Brij76-formulation leaked at 37°C (24.9%, 30 min) (Fig. 4C). In contrast, the Brij78 formulation (HaT) displayed full drug release within 3 min at both 40 and 42°C, and no release could be detected at 30 and 37°C (Fig. 4D). The Brij700 formulation had a similar release profile to the Brij78-formulation at 40 and 42°C, but leaked substantially at 37°C (22.5% at 30 min) (Fig. 4E). Measurement of the DOX release from the Brij98-formulation was also performed, but no drug release could be detected (data not shown). The Brij78-formulation was also loaded with GEM and CDDP, and a similar set of release profiles were generated, both of which appear similar to the DOX release profile (Fig. 5).

Example 4: The influence of the mol% of Brii78 on the thermal sensitivity of the formulation

To investigate the influence of Brij78 on the thermal sensitivity of the formulation, we determined the drug release profile of the Brij78-liposomes containing 1-8 mol% of Brij 78 in comparison with the DPPC-liposomes (0% Brij 78) (Fig. 6). The DPPC- liposomes which did not contain Brij78 showed significantly slower drug release at 42°C compared to the Brij78-liposomes (1-8 mol%) (Fig. 6A). As the mol% of Brij78 rose, the observed release profile also increased, with a plateau reached from 4-8 mol%. While the DPPC-liposomes exhibited DOX leakage at 37°C (Fig. 6B) (31.9%, 30 min), increasing levels of Brij 78 content improved the stability of the formulation at 37°C, with a maximum level of 3% leakage measured from liposomes containing 4-8 mol% Brij 78.

Example 5: The stability of HaT formulation in storage

There was no diameter change in the HaT and LTSL formulations during one month of storage at 4°C (Table 2) with very little drug leakage (<3%).

Table 2. Stability of the formulations in storage at 4°C in the dark

101.4 101.7 101.6 101.8

HaT Diameter (±0.5) (±0.4) (±0.5) (±0.6)

Drug content remaining 100 99.56 98.71 97.49 in the liposomes (%) (±0.32) (±0.19) (±0.09) (±0.19)

Example 6: Temperature dependent release of the encapsulated content from liposomes measured by MRI

The Tl map images of the formulations after heated at 37-42°C for 3 min are shown in Fig. 7 and the Tl values of the samples are summarized in Table 3. The Tl signal was greatly attenuated when the HaT formulation was heated at 40-42°C, while the effect was only significant for LTSL when heated above 42°C. The DPPC-liposomes displayed relatively smaller changes at 40-42°C compared to HaT and LTSL. The Triton-X 100 treated samples were included as a positive control of complete release of the encapsulated content, which exhibited the most significant reduction of the signal.

Table 3. Tl values (ms) of the formulations after different treatments measured by MRI

Example 7: Effect of post-inserting Brij78

into the liposomes on the release of drug

We compared two different preparation methods for the incorporation of Brij 78: the post-insertion method and the conventional thin film hydration method. Both techniques resulted in liposomes exhibiting a temperature dependent drug release profile (Fig. 8). As shown in Fig. 8, post-insertion of 8 mol% Brij78 into preformed DPPC liposomes yielded a formulation with a DOX release profile similar to the release profile of conventionally prepared DPPC/Brij 78 (96/4) formulation, with 100% release at 42°C in 3 min and little leakage at 37°C. Example 8: Intracellular uptake and cytotoxicity of DOX in different formulations

Cells were cultured with DOX in different formulations (which had been pre-heated to stimulate DOX release), and the intracellular uptake of DOX was measured quantitatively. As shown in Fig. 9A, the cells treated with the HaT formulation preheated at 40°C or 42°C showed similar intracellular uptake of DOX compared to the cells treated with the same concentration of free DOX. In contrast, the cells treated with the HaT formulation at 37°C showed significantly reduced intracellular uptake of DOX. In comparison, the cells treated with the LTSL formulation preheated at 42°C also displayed similar DOX uptake compared to free DOX, but the cells treated with the LTSL preheated at 37°C or 40°C did not display a significantly increased intracellular uptake of DOX. Fluorescent microscopy images were collected to support the quantitative data, as DOX accumulates in the nucleus (Fig. 9B-D).

The intracellular uptake of DOX was further confirmed with the cytotoxicity analysis (Fig. 10). The HaT formulation preheated at 40°C or 42°C induced a cytotoxicity profile similar to that of free DOX. In comparison, the LTSL formulation preheated at 42°C induced similar cytotoxicity compared to free DOX, but the LTSL formulation preheated at 40°C did not induce significant cytotoxicity. Both liposomal formulations were only slightly toxic when incubated at 37°C. The cytotoxicity of free DOX was not dependent on temperature. The liposomal carriers alone showed little cytotoxicity for both HaT & LTSL (Data not shown). The IC50 values are summarized in Table 4.

Table 4. IC₅₀ value

Example 9; Blood PK

The blood PK of DOX in different formulations was investigated in tumor-free mice (Fig. 11). A similar PK profile was observed for the LTSL and HaT formulations. Approximately 50 % of the injected dose (ID) remained in the blood circulation at 30 min after the administration of LTSL or HaT, and 40% remained after 1 h. In comparison, only 10% of the ID was recovered in the plasma of the mice treated with free DOX after 10 min, and no DOX could be detected after 1 h. The AUC (area under the curve calculated from time of injection to 1 h) was analyzed for each mouse; the AUC of the HaT formulation was similar to LTSL (2.5 and 2.7 mg h/ml, respectively), and these values were 6.5 times higher compared to the AUC of the mice receiving free DOX (0.38 mg h/ml). Example 10: Drug release profiles and PK of DOX-loaded liposomes prepared by copper ion gradient method

HaT(Cu) DOX liposomes prepared using a Cu²⁺ gradient exhibited >95% drug loading, but release of DOX from the 4 mol% Brij 78 liposome composition at 37-42°C was poor compared to the HaT liposomes prepared using a citric acid pH gradient (Figure 11.1 A-C). Post-insertion of Brij 78 in preformed HaT(Cu) liposomes with 0-32 mol% Brij 78 was performed, and the new liposomes were analyzed by dynamic light scattering to ensure the absence of Brij 78 micelles. HaT(Cu) DOX liposomes post- inserted with 24 mol% Brij 78 exhibited a drug release profile similar to HaT(citric acid) (Figure 11.1 A-C), drug leakage at 37°C in serum-containing solutions was significantly reduced (Figure 11.1 D), and blood PK was improved compared to HaT(citric acid) and LTSL (Figure 11.1 E).

Example 11: Preliminary toxicity study

The body weight of the tumor-free mice was tracked after administration of different DOX formulations (Fig. 12). There were no significant differences among the three groups when 3 mg DOX/kg was administered. Dose escalation for the HaT and LTSL to 10 mg DOX/kg also caused no effect. However, the mice treated with 10 mg/kg of free DOX experienced significant weight loss by day 7. The blood compatibility of the carriers was investigated by the hemolysis assay (Fig. 13). The LTSL and HaT formulations displayed similar blood compatibility.

Example 12: Biodistribution study

The distribution of DOX to the tumor and other normal tissues 1 h subsequent to the i.v. injection of DOX in different formulations in combination with local hyperthermia (42°C) was studied (Fig. 14). Treatment of the mice with HaT and LTSL led to significantly increased DOX uptake in the heated tumor compared to free DOX solution, by 5.2-fold and 3.8-fold, respectively. HaT enhanced the delivery of DOX to the heated tumor by 1.4-fold compared to LTSL. The amount of DOX in the tumor treated with hyperthermia combining with the liposomal formulations was significantly higher compared to that in the unheated tumor (HaT, 6.6-fold; LTSL, 2.8-fold). There was no significant difference in DOX accumulation between the heat-treated and unheated tumors in the group receiving free DOX. With regards to blood concentration, 7.2% of ID/g DOX was present in plasma at 1 h after the injection of LTSL, whereas DOX was almost completely eliminated from the blood for the free DOX and HaT group (0.6% ID/g and 2.8%ID/g, respectively). DOX concentrations in the lung, kidney, liver and heart were significantly reduced for the HaT and LTSL groups compared to free DOX by 1.3- to 15-fold, with a 10-fold decreased heart accumulation. The ratios of DOX content in tumor/heart were 0.3 ± 0.1 (free DOX), 15.6 ± 6.4 (LTSL) and 24.6 ± 3.3 (HaT).

Example 13: In vivo anticancer study

As seen in the histology images (Fig. 15A-C), the tumor mass was apparent in all unheated treatments. The free DOX treated tumor mass was unaffected in both legs (Fig. 15A), whereas the LTSL/heat treated tumor was substantially inhibited (Fig. 15B) and no tumor cells could be detected in the HaT/heat treated tumor (Fig. 15C). The tumors in the mice treated with free DOX showed initial growth inhibition (Fig. 15D), but at later time points, tumor growth resumed. In comparison, the tumors in the mice treated with hyperthermia and the HaT and LTSL formulations were significantly inhibited at all time-points measured, and two HaT-treated mice and one LTSL-treated mouse experienced complete remission. The unheated tumors treated with different formulations did not display any significant growth reductions compared to the untreated tumor. No decrease in body weight was observed for any of the mice treated (data not shown). Hyperthermia alone did not show any effect in slowing tumor growth. Referring to Figure 17, a decline in the tumor volume is detected for 2/4 LTSL subjects and 4/4 HaT subjects (heated left leg compared to control right leg) after a single and low dose treatment (3 mg DOX/kg), with complete elimination of tumor in 2/4 HaT mice. The mice that experience complete regression of the primary tumor (1/4 for LTSL and 2/4 for HaT) also display complete remission of the distal tumor 15-25 days later. Example 14: Effect of HaT formulation on ADR cells

The DOX, LTSL and HaT formulations all caused similar decreases in A2780 (drug sensitive) cell viability (Fig. 16A), with -80% inhibition at 3 μΜ DOX. In contrast, the A2780-ADR cells were significantly less responsive to DOX treatment, with the exception of the group treated with HaT, which exhibited a 95% loss of viability at 10 μΜ DOX (Fig. 16B).

Example 15: Characteristics of DOX-loaded DPPC Liposomes

with various Brii and Myri surfactants

The thermosensitive characteristics of a range of surfactants in DPPC liposomes were determined in accordance with the described methods. Referring to Figure 18, the liposomes exhibiting advantageous drug loading efficiency, low leakage at 37°C, and high release at 42°C were submitted to a structure-activity relationship (SAR) analysis. Detailed analysis of selected surfactants with proton NMR and MS was carried out to determine chemical composition and molecular weight. The molecular formula for each surfactant in Table 5 is reported by the supplier, but importantly, there is a distribution in the molecular weights of the PEG blocks, and the molecular formula therefore represents only an average composition. Figure 19 depicts the ES^" MS spectra for Brij 78, demonstrating a typical molecular weight distribution for these polyethoxylated compounds. Table 5 below summarizes the chemical composition analysis, including the estimated PEG repeat units based on NMR and MS analysis, and the linker chemistry between the C16/18 chain and the PEG chain (ester or ether).

Table 5: Composition analysis of Brij and Myrj surfactants that yield thermosensitive properties when formulated with DPPC at 4 mol%.

Figure 20 further depicts the relationship between composition and thermosensitive function for molecules defined by the composition Ci8H3₇(OCH₂CH₂)_nOH, wherein n = 2-100.

Example 16: Characteristics of DOX-loaded Brij 78 Liposomes

with different phospholipids The characteristics of Brij 78 Liposomes with different phospholipids were determined in accordance with the described methods. Referring to Figure 18, only the DPPC/Brij78 liposome was shown to have all the advantageous characteristics of high drug loading efficiency, low leakage at 37°C and high release at 42°C. Example 17: Synthesis of Brii 78-cRGDPK peptide conjugate

To a solution of succinic anhydride (0.521 g, 5.21 mmol) in DMF with DMAP (0.021 g, 0.174 mmol) and TEA (0.018 g, 0.174 mmol) was added Brij 78 (2 g, 1.735 mmol) in a solution in DMF. The reaction was stirred under nitrogen for 12 hours, extracted with ether, washed with IN HC1 and water, dried over magnesium sulphate and concentrated. The extracted residue was purified by silica gel column chromatography to yield Brij 78-acid (Fig 21): H NMR (CDC1₃): 4.26 (succinic CH₂), 3.62 (CH₂-0), 2.64 (succinic CH₂), 1.27 (aliphatic CH₂), 0.87 (CH₃). LC/MS (CI 8 column, water/acetonitrile gradient, 5%-95% over 20 min): 13.6 min, m/z = 721 - 1514. The Brij 78-acid (0.450g, 0.36 mmol) was dissolved in DMF, and to this was added EDC HC1 (0.137g, 0.720 mmol), NHS (0.083g, 0.72 mmol), and TEA (0.004g, 0.036 mmol). A cyclic peptide (cRGDPK) was dissolved in DMF, added to the reactor, and the solution was stirred overnight. The DMF was removed by rotoevaporation, and the crude product was analyzed by LC/MS: 11.64 min, m/z = 1617 - 1910. Mass of product (Fig 21) corresponds to the mass of Brij 78-acid reacted with cRGDPK (m/z = 602).

Example 18: Release of GEM from HaT and LTSL formulations in FBS- containing solution Both GEM HaT and LTSL HaT formulations are stable at 37-38°C with no drug leakage in serum for 30 min (Fig. 22). Approximately 20%, 80%, 95% and 100% of drug release is detected from HaT-GEM at 39, 40, 41 and 42°C in 2 min, respectively. The release rate of GEM from LTSL is significantly slower, with about 0, 0, 20 and 95% release at 39, 40, 41 and 42°C in 2 min, respectively. The improved release profile of GEM HaT is further exemplified by comparison of the drug release rate constants, which are significantly higher for HaT-GEM than LTSL-GEM by 28-, 8- and 1.6-fold, at 40, 41 and 42°C, respectively. Example 19; HaT-GEM PK and biodistribution

Blood clearance of GEM is reduced by 50-fold relative to free GEM for both of the liposomal formulations (Fig. 23). More importantly, almost all (>95%) of the administered dose remains in the blood circulation for the temperature-sensitive liposomes (TSL) formulations during the first hour when the hyperthermia treatment would be applied, indicating the stability of the TSL formulations is optimal. We have also compared the drug delivery of free GEM, HaT-GEM and LTSL-GEM, to the heated and unheated tumor in a murine model; wherein the mice had two tumors each, one of which was locally heated at 43°C for 1 h after the i.v. injection of the formulations. Drug uptake in the unheated tumor is comparably low for all three formulations, but the GEM delivery to the heated tumor by HaT is increased by 25-fold and 7-fold compared to that of free GEM and LTSL, respectively. Interestingly, the remaining dose in the blood for LTSL is 3-fold higher than that for HaT, confirming the faster and more complete drug release for the HaT formulation at hyperthermic temperatures. In fact, the drug delivery to the heated tumor for LTSL-GEM is not improved relative to free GEM, because of its slow release kinetics, and instead the majority of the dose remains in the blood.

Example 20: Efficacy study of GEM HaT in Mice Bearing Footpad Tumors In the footpad s.c. tumor models (PAN02 and LL/2), a single dose of HaT-GEM at 20 mg/kg in combination with localized footpad heating with a warm water bath at 43 °C completely regresses the tumors in 5 days, while LTSL-GEM and free GEM only display little to modest activity (Figure 24). The regressed tumors were monitored for 15-20 days after remission, confirming cure. It is noted that these three formulations are compared at the MTD (20 mg GEM/kg for liposomal GEM) and the maximum deliverable dose (120 mg GEM/kg for free GEM) for one single i.v. dose. Treatment with hyperthermia or liposomal formulations alone does not exhibit significant efficacy. Prophetic Example

The Brij 78-cRGDPK peptide conjugate of Example 17 is purified by preparative scale LC/MS. The lipid is used to prepare RGD coupled HaT nanoparticles for delivery of DOX to tumors. Brij 78-cRGDPK (1 mol%) and Brij 78 (3 mol%) are mixed with DPPC (96 mol%) in IPA, a thin film is prepared, hydrated with 300 mM citric acid, and sized by extrusion. Liposomes are loaded with DOX by the pH gradient method (1:5 DOX:lipid), and are purified by gel filtration with HBS. Liposome concentration is adjusted and BALB/c EMT-6 foot tumor-bearing mice are injected with 3 mg/kg DOX doses. The tumor is heated to 42°C for 1 hour after i.v. injection, after which the mice are sacrificed and a biodistribution analysis of DOX is conducted. Drug delivery between the control HaT and RGD-HaT is compared, and it is expected that RGD-HaT will exhibit improved drug delivery.

DISCUSSION Liposome technologies for the encapsulation and delivery of therapeutic drugs have been developed with the objective to enhance therapy and minimize side effects. To improve the PK, the introduction of the long circulating liposomal formulation (HSPC/Cholesterol/PEG₂ooo-DSPE) has been successful, and is in clinical application (Doxil®) [2]. However, the long-circulating Doxil^® liposome does not appear to release the drug effectively in the tumor and generates new side effects including hand-foot syndrome due to the much prolonged PK [5], therein undermining the benefits [6]. To address this shortfall, a variety of liposomal formulations that can be triggered to release the drug under specific mechanisms applied in the tumor have been developed [11] [79]. Thermosensitive liposomes were first introduced by Yatvin et al. [31] and Weinstein et al. [32], wherein they formulated drug into DPPC-liposomes: DPPC has a membrane transition temperature of 42°C, and accordingly, the drug could be triggered to release in the heated tumor (42°C) for improved drug delivery. Other lipid components, such as DSPC, cholesterol or DSPE-PEG were later added to stabilize the formulation in the blood [30, 44, 47, 51, 55, 80-81], however, these components reduced thermal sensitivity, leading to similar drug delivery compared to free DOX[44, 47]. Needham and Dewhirst optimized the formulation by introducing lyso-PC and DSPE-PEG into the DPPC-liposomes [19-21, 28-29, 51, 64, 82-85], which significantly enhanced its stability at 37°C but improved the thermal sensitivity at the same time [19-21, 28-29, 51, 64, 82-85].

The mechanism of drug release from LTSL has been investigated by several groups [60, 86-88]. The current model states that the lyso-PC and DSPE-PEG (both are detergent-like molecules) are enriched in the boundary of the DPPC lipid rafts of the LTSL liposomes, decreasing the transition temperature at the boundary (<42°C) [88]. According to this model, when heated at 39-41°C, the membrane becomes permeable at the boundaries, resulting in rapid release of the drug [88]. Mills et al. measured the ion permeability and established that it was consistent with the observed drug release profile [61]. There remains however no clear consensus on the role of each LTSL component, as mechanism studies are not consistent in their conclusions: for example, Li et al prepared liposome composed of DPPC/DSPC/PEG₂ooo-DSPE without lysolipid [89], and showed that increasing PEG density facilitated the release of drug from liposomes, but Banno et al reported that increased PEG content did not affect release rates [90]. The membrane model for LTSL suggests that the presence of lyso-PC and DSPE-PEG in both the leaflets of the lipid bilayer is necessary for forming the boundary for drug release. To further elucidate if the HaT formulation relied on a similar release mechanism, we prepared the HaT liposomes using two different methods: the post- insertion method and the conventional thin film method (Fig. 8). The post-insertion method was employed to incorporate Brij 78 only onto the outer leaflet of the liposomes. Nevertheless, the resulting formulation (DPPC-liposomes post-inserted with 8 mol% Brij 78) displayed similar release kinetics compared to the liposomes prepared with the thin film hydration method (DPPC/Brij 78 = 96/4) (Fig. 8). The data suggest that the HaT formulation might exhibit a distinctive mechanism for drug release. Direct comparisons of drug release in response to hyperthermia conditions were drawn between the HaT and LTSL formulations, to determine if the incorporation of the Brij- surfactant in the membrane would influence release profiles. The release of drug from the LTSL is well characterized: for example, Mills et al. [61] showed that DOX was completely released from LTSL within 5 min at 42°C in vitro, and Woo et al. [63] demonstrated that CDDP was released from LTSL in 5 min at 42°C. In both reports, drug did not leak from the LTSL at 37°C. However, Mills et al. reported that LTSL released only 40-50% of DOX at 40°C [61]: we have confirmed the above release characteristics of LTSL in our study (Fig. 4B). In comparison, the novel HaT liposomal formulation released 100% of DOX within 3 min at 40°C, the DOX released from the HaT was taken up into the cells efficiently (Fig. 9), and induced a cytotoxicity similar to that free DOX (Fig. 10). Similarly, 100% of GEM and 80-100% of CDDP were released from the HaT formulation within 2 min at 40°C (Fig. 5A and B). More importantly, little drug was detected leaking out from HaT at 37°C. These results confirm that the HaT formulation is an improvement over LTSL.

The data also show that HaT prepared with the Cu²⁺ gradient method followed by post- insertion of 24 mol% Brij78 [HaT(Cu-24%)] exhibited similar temperature-sensitive release profile compared to HaT prepared with the pH gradient method [HaT(Cit)]. The stability of HaT(Cu-24%) was significantly enhanced in the serum containing medium at 37°C, which corresponds to the significantly improved pharmacokinetics in the mouse.

The range of surfactant compositions described in Figure 18 includes compounds which do not perform in this application, and includes surfactants with low PEG content and/or unsaturated chemistry. From the data summarized in Table 5, an SAR analysis to link chemical composition and functional performance was performed. The surfactant structures defined as:

(1) C_{l 7}H3₅(CH₂)p(CO)_q(OCH₂CH₂)_nOH, wherein n = 10-100; p = 0 or 1; q - 0 or 1 ; and p+q = 1 ; and

(2) Ci₆H₃₃(OCH₂CH₂),oOH (Brij 56) impart thermosensitive properties to DPPC liposome formulations.

The Brij/Myrj surfactants tested each represent a range of molecular weights, not a discrete molecular identity (Figure 19). The relationship between composition and thermosensitive function for molecules defined by the composition Ci₇H₃5(CH₂)_p(CO)_q(OCH₂CH₂)_nOH, wherein n = 2-100; p = 1; and q = 0 (i.e. C₁₈H₃₇(OCH₂CH₂)_nOH, wherein n = 2-100) indicates that at n = 5 PEG repeat units (NMR estimation), thermosensitive properties are introduced (Figure 20).

Clearly, the data indicates that thermosensitive properties conferred by polyethoxylated surfactants are restricted to a specific set of compositions. Achieving homogeneous temperature distribution in tumors using currently available heating technologies has been a challenge. For example, Brown et al. reported that a thermal gradient greater than 1°C /mm was observed in the KHT fibrosarcoma tumor when it was immersed in a water bath [91]. Difficulty with thermal control using radiofrequency hyperthermia has been reported, with significant intratumoral temperature gradients and variation noted between patients [92]. Similarly, it has been reported that MRgFUS induced a thermal gradient with heated tissue [93]. Therefore, designing a novel carrier that exhibits an improved thermal sensitivity (a more rapid kinetic of release at the lower range of hyperthermia, i.e. 39-40°C) is expected to enhance drug delivery with heterogeneously heated tumor environments. In the mouse PK study, more than 40% ID was recovered for the liposomal DOX (LTSL-DOX and HaT-DOX) 1 h after the injection in tumor-free mice, while the concentration of free DOX was barely detectable (Fig. 11), suggesting that both the HaT and LTSL formulations could retain the drug in the blood of the mice that were under no hyperthermia. However, after 1 h of treatment with hyperthermia at the tumor (42°C), the DOX remaining in the blood was found to be 0.3%, 7.2% and 2.8% for free DOX, LTSL-DOX and HaT-DOX, respectively (based on the assumption of 1 ml of plasma per mouse) (Fig. 14). As demonstrated in Fig. 11, only free DOX was eliminated quickly from the blood within 1 h; and therefore, the biodistribution data suggest that the drug release from HaT was significantly more complete for HaT than for LTSL, and this observation is further confirmed by a measured 1.4-fold increase in tumor DOX content for the HaT formulation (Fig. 14). The improved PK and biodistribution results for the HaT group were likely due to the heterogeneous thermal distribution in the heated tumor, which favored rapid release from HaT formulation at the lower range of the hyperthermic temperatures (i.e. 40°C). The positive PK and biodistribution results for the HaT tests were also supported by the measurement of significantly enhanced antitumor effect (Fig. 15). These results suggest that this novel formulation can compensate for the hurdles presented by tumor heterogeneity and may be useful to further improve the tumor delivery compared to the current LTSL formulation. Furthermore, complete remission of the distal tumor (15-25 days later) in mice that experience complete regression of the primary tumor (Fig. 17) potentially suggest additional immune stimulation by HaT-DOX.

In the treatment of cancer with DOX, cardiotoxicity is a serious clinical limitation [94], and accordingly, this study examined the biodistribution of the DOX administered as free drug, and as liposomal formulation. To our knowledge, this is the first study examining the comprehensive biodistribution of DOX from thermosensitive liposomes in combination with mild hyperthermia. In the biodistribution study, the accumulation of DOX from both liposomal formulation (LTSL or HaT) in the heart was found to be significantly lower compared to free DOX (6.7% ID/g, 0.46% ID/g and 0.57% ID/g for free DOX, HaT and LTSL, respectively) (Fig. 14), clearly indicating that both liposomal formulations were safer than free DOX. In addition, it is well known that liposomes are taken up by the MPS in the liver and spleen [2] [3]. However, the DOX accumulation in the liver and spleen of mice treated with liposomal formulation was either comparable or >2-fold lower than those treated with free DOX (Fig. 14). These results confirm that the majority of DOX was released in the heated tumor, leaving little remaining DOX or liposomal DOX to accumulate within the MPS or other normal tissues.

In a further toxicity study, the identity of Brij78 as a surfactant and a possible mediator of membrane lysis was considered: a hemolysis assay (Fig. 13) confirmed that the HaT displayed similar blood compatibility compared to the LTSL formulation, suggesting Brij78 was tightly associated with the liposomes and therefore, did not induce membrane lysis. Gadolinium has been previously studied as an indicator of drug release from the liposomal formulations in vitro and in vivo by MRI [62, 84, 95]. Similarly, Mn²⁺ and DOX were co-encapsulated into the LTSL-liposomes for the study of drug release in the tumor [84]. The drug delivery and the MRI signal attenuation was found to correlate well [84] [96] [97]. However, Mn²⁺ is toxic and has not been approved for clinical use. Gd³⁺ and Mn²⁺ produce MRI contrast only when water can freely exchange with the molecule: when encapsulated, the exchange is inhibited, and the MRI contrast is not significant. Accordingly, when Gd³⁺ or Mn²⁺ are released into the aqueous environment surrounding the liposomes, MRI contrast is detected. In this study, we investigated the release of the encapsulated content from the HaT formulation at different temperatures by MRI (Fig. 7): the results are consistent with that of the drug release profile (Fig. 4). The HaT-Gd formulation might be used for monitoring the drug release/delivery in the tumor using non-invasive MRI. As the interior phase (Gd- DTPA) of the liposomal formulation had a pH around 4, DOX was remote-loaded into the Gd³⁺ containing HaT-liposomes with high efficiency (drug/lipid = 1/20, >95% loading) (data not shown). The incorporation of DTPA therefore provided the pH gradient enabling drug loading in the liposomes, and reduced the toxicity of Gd³⁺ [98].

It has been reported that Brij78 affects p-glycoprotein function, and accordingly, cells exposed to the Brij78-containing nanoemulsion formulation lost their MDR phenotype [99]. As HaT is formulated with 4 mol% Brij78, it was considered that DOX delivered with HaT-liposomes may be effective against MDR tumor cells. Indeed, the data from this study confirmed that the HaT formulation had the effect of enhancing DOX action on the resistant clone.

A HaT liposome labelled with a ligand such a targeting moiety recognizing antigens on the tumor vasculature may be expected to display increased retention in the tumor microvaculature, and accordingly, increase the dose released within the locally heated tumor. An example of this concept is an RGD-labelled HaT liposome. The RGD- labelled HaT is expected to bind to tumor cells overexpressing integrins, as the RGD peptide binds to certain classes of these integrins. The selective slowing or arrested movement of the RGD-HaT liposomes through the heated tumor is expected to reduce the transport of drug and liposome out of the tumor, leading to increased local drug release within the tumor and increased action against the tumor cells.

In conclusion, the HaT formulation was designed to release the drug under mild hyperthermia generated locally in the targeted tumor: in vitro and in vivo data indicate that HaT could be stimulated to release the drug, and the cells and the tumor were responsive to the improved delivery of the drug. Furthermore, the rapid drug release at the lower temperature (40°C) offered the HaT formulation an advantage over LTST in delivering an increased amount of the drug to the heterogeneously heated tumor. In addition, the replacement of DSPC-PEG and MSPC with Brij78 not only conferred both stealth and thermosensitivity properties, but it also assisted in overcoming drug resistance. This simple formulation might also offer advantages on manufacturing, scale-up, and costs, as it is prepared using fewer and less expensive materials.

Although preferred embodiments of the invention have been described herein, it will be understood by those skilled in the art that variations may be made thereto without departing from the spirit of the invention or the scope of the appended claims. All references mentioned herein, including in the following list of references, are incorporated by reference in their entirety.

Reference List

1. Allen, T.M. and P.R. Cullis, Drug delivery systems: entering the mainstream.

Science, 2004. 303(5665): p. 1818-22.

2. Gabizon, A., et al., Prolonged circulation time and enhanced accumulation in malignant exudates of doxorubicin encapsulated in polyethylene-glycol coated liposomes. Cancer Res, 1994. 54(4): p. 987-92.

3. Li, S.D. and L. Huang, Pharmacokinetics and biodistribution of nanoparticles.

Mol Pharm, 2008. 5(4): p. 496-504.

4. Matsumura, Y. and H. Maeda, A new concept for macromolecular therapeutics in cancer chemotherapy: mechanism of tumoritropic accumulation of proteins and the antitumor agent smancs. Cancer Res, 1986. 46(12 Pt 1): p. 6387-92.

5. Plosker, G.L., Pegylated liposomal Doxorubicin: a review of its use in the

treatment of relapsed or refractory multiple myeloma. Drugs, 2008. 68(17): p. 2535-51.

6. Gabizon, A., H. Shmeeda, and Y. Barenholz, Pharmacokinetics of pegylated liposomal Doxorubicin: review of animal and human studies. Clin

Pharmacokinet, 2003. 42(5): p. 419-36.

7. Laginha, K.M., et al., Determination of doxorubicin levels in whole tumor and tumor nuclei in murine breast cancer tumors. Clin Cancer Res, 2005. 11(19 Pt 1): p. 6944-9.

8. Harrington, K.J., et al., Phase I-II study of pegylated liposomal cisplatin (SPI- 077) in patients with inoperable head and neck cancer. Ann Oncol, 2001. 12(4): p. 493-6. Bandak, S., et al., Pharmacological studies of cisplatin encapsulated in long- circulating liposomes in mouse tumor models. Anticancer Drugs, 1999. 10(10): p. 911-20.

Maeda, H., Tumor-selective delivery of macromolecular drugs via the EPR effect: background and future prospects. Bioconjug Chem, 2010. 21(5): p. 797- 802.

Andresen, T.L., S.S. Jensen, and K. Jorgensen, Advanced strategies in liposomal cancer therapy: problems and prospects of active and tumor specific drug release. Prog Lipid Res, 2005. 44(1): p. 68-97.

Lindner, L.H. and M. Hossann, Factors affecting drug release from liposomes. Curr Opin Drug Discov Devel. 13(1): p. 111-23.

Meng, F., Z. Zhong, and J. Feijen, Stimuli-responsive polymersomes for programmed drug delivery. Biomacromolecules, 2009. 10(2): p. 197-209.

Gerasimov, ON., et al., Cytosolic drug delivery using pH- and light-sensitive liposomes. Adv Drug Deliv Rev, 1999. 38(3): p. 317-338.

Shum, P., J.M. Kim, and D.H. Thompson, Phototriggering of liposomal drug delivery systems. Adv Drug Deliv Rev, 2001. 53(3): p. 273-84.

Huang, S.L. and R.C. MacDonald, Acoustically active liposomes for drug encapsulation and ultrasound-triggered release. Biochim Biophys Acta, 2004. 1665(1-2): p. 134-41.

Schroeder, A., J. Kost, and Y. Barenholz, Ultrasound, liposomes, and drug delivery: principles for using ultrasound to control the release of drugs from liposomes. Chem Phys Lipids, 2009. 162(1-2): p. 1-16.

Meers, P., Enzyme-activated targeting of liposomes. Adv Drug Deliv Rev, 2001 53(3): p. 265-72. 19. Hauck, M.L., et al., Phase I trial of doxorubicin-containing low temperature sensitive liposomes in spontaneous canine tumors. Clin Cancer Res, 2006. 12(13): p. 4004-10.

20. Needham, D. and M.W. Dewhirst, The development and testing of a new

temperature-sensitive drug delivery system for the treatment of solid tumors.

Adv Drug Deliv Rev, 2001. 53(3): p. 285-305.

21. Kong, G., et al., Efficacy of liposomes and hyperthermia in a human tumor xenograft model: importance of triggered drug release. Cancer Res, 2000. 60(24): p. 6950-7. 22. Poon, R.T. and N. Borys, Lyso-thermosensitive liposomal doxorubicin: a novel approach to enhance efficacy of thermal ablation of liver cancer. Expert Opin Pharmacother, 2009. 10(2): p. 333-43.

23. Davies Cde, L., et al., Radiation improves the distribution and uptake of

liposomal doxorubicin (caelyx) in human osteosarcoma xenografts. Cancer Res, 2004. 64(2): p. 547-53.

24. Gardner, R.A., et al., Focused microwave phased array thermotherapy for primary breast cancer. Ann Surg Oncol, 2002. 9(4): p. 326-32.

25. Dromi, S., et al., Pulsed-high intensity focused ultrasound and low temperature- sensitive liposomes for enhanced targeted drug delivery and antitumor effect. Clin Cancer Res, 2007. 13(9): p. 2722-7.

26. Salomir, R., et al., Local hyperthermia with MR-guided focused ultrasound: spiral trajectory of the focal point optimized for temperature uniformity in the target region. J Magn Reson Imaging, 2000. 12(4): p. 571-83.

27. Morawski, A.M., G.A. Lanza, and S.A. Wickline, Targeted contrast agents for magnetic resonance imaging and ultrasound. Curr Opin Biotechnol, 2005.

16(1): p. 89-92. 28. Needham, D., et al., A new temperature-sensitive liposome for use with mild hyperthermia: characterization and testing in a human tumor xenograft model. Cancer Res, 2000. 60(5): p. 1197-201.

29. Yarmolenko, P.S., et al., Comparative effects of thermosensitive doxorubicin- containing liposomes and hyperthermia in human and murine tumours. Int J

Hyperthermia. 26(5): p. 485-98.

30. Morita, K., et al., Efficacy of doxorubicin thermosensitive liposomes (40

degrees C) and local hyperthermia on rat rhabdomyosarcoma. Oncol Rep, 2008. 20(2): p. 365-72. 31. Yatvin, M.B., et al., Design of liposomes for enhanced local release of drugs by hyperthermia. Science, 1978. 202(4374): p. 1290-3.

32. Weinstein, J.N., et al., Liposomes and local hyperthermia: selective delivery of methotrexate to heated tumors. Science, 1979. 204(4389): p. 188-91.

33. Yatvin, M.B., et al., Selective delivery ofliposome-associated cis- dichlorodiammineplatinum(II) by heat and its influence on tumor drug uptake and growth. Cancer Res, 1981. 41(5): p. 1602-7.

34. Tacker, J.R. and R.U. Anderson, Delivery of antitumor drug to bladder cancer by use of phase transition liposomes and hyperthermia. J Urol, 1982. 127(6): p. 1211-4. 35. Maekawa, S., K. Sugimachi, and M. Kitamura, Selective treatment of metastatic lymph nodes with combination of local hyperthermia and temperature-sensitive liposomes containing bleomycin. Cancer Treat Rep, 1987. 71(11): p. 1053-9.

36. Bassett, J.B., et al., Treatment of experimental bladder cancer with

hyperthermia and phase transition liposomes containing methotrexate. J Urol, 1988. 139(3): p. 634-6. 37. Khoobehi, B., et al., Hyperthermia and temperature-sensitive liposomes:

selective delivery of drugs into the eye. Jpn J Ophthalmol, 1989. 33(4): p. 405- 12.

38. Zou, Y.Y., et al., Targeting behavior of hepatic artery injected temperature sensitive liposomal adriamycin on tumor-bearing rats. Sel Cancer Ther, 1990.

6(3): p. 119-27.

39. Nishimura, Y., et al., Treatment of murine SCC VII tumors with localized

hyperthermia and temperature-sensitive liposomes containing cisplatin. Radiat Res, 1990. 122(2): p. 161-7. 40. Merlin, J.L., Encapsulation of doxorubicin in thermosensitive small unilamellar vesicle liposomes. Eur J Cancer, 1991. 27(8): p. 1026-30.

41. Iga, K., et al., Increased tumor cisplatin levels in heated tumors in mice after administration of thermosensitive, large unilamellar vesicles encapsulating cisplatin. J Pharm Sci, 1991. 80(6): p. 522-5. 42. Iga, K., Y. Ogawa, and H. Toguchi, Heat-induced drug release rate and

maximal targeting index of thermosensitive liposome in tumor-bearing mice. Pharm Res, 1992. 9(5): p. 658-62.

43. Zou, Y., et al., Enhanced therapeutic effect against liver W256 carcinosarcoma with temperature-sensitive liposomal adriamycin administered into the hepatic artery. Cancer Res, 1993. 53(13): p. 3046-51.

44. Maruyama, K., et al., Enhanced delivery of doxorubicin to tumor by long- circulating thermosensitive liposomes and local hyperthermia. Biochim Biophys Acta, 1993. 1149(2): p. 209-16.

45. Huang, S.K., et al., Liposomes and hyperthermia in mice: increased tumor uptake and therapeutic efficacy of doxorubicin in sterically stabilized liposomes. Cancer Res, 1994. 54(8): p. 2186-91. Ning, S., et al., Hyperthermia induces doxorubicin release from long- circulating liposomes and enhances their anti-tumor efficacy. Int J Radiat Oncol Biol Phys, 1994. 29(4): p. 827-34. Unezaki, S., et al., Enhanced delivery and antitumor activity of doxorubicin using long-circulating thermosensitive liposomes containing amphipathic polyethylene glycol in combination with local hyperthermia. Pharm Res, 1994. 11(8): p. 1180-5. Chelvi, T.P., S.K. Jain, and R. Ralhan, Hyperthermia-mediated targeted delivery of thermosensitive liposome-encapsulated melphalan in murine tumors. Oncol Res, 1995. 7(7-8): p. 393-8. Kakinuma, K., et al., Drug delivery to the brain using thermosensitive liposome and local hyperthermia. Int J Hyperthermia, 1996. 12(1): p. 157-65. Kakinuma, K., et al., Targeting chemotherapy for malignant brain tumor using thermosensitive liposome and localized hyperthermia. J Neurosurg, 1996. 84(2): p. 180-4. Gaber, M.H., et al., Thermosensitive liposomes: extravasation and release of contents in tumor microvascular networks. Int J Radiat Oncol Biol Phys, 1996. 36(5): p. 1177-87. Chelvi, T.P. and R. Ralhan, Hyperthermia potentiates antitumor effect of thermosensitive-liposome-encapsulated melphalan and radiation in murine melanoma. Tumour Biol, 1997. 18(4): p. 250-60.

Sharma, D., et al., Thermosensitive liposomal taxol formulation: heat-mediated targeted drug delivery in murine melanoma. Melanoma Res, 1998. 8(3): p. 240- 4. 54. Nishita, T., Heat-sensitive liposomes containing cisplatin and localized hyperthermia in treatment of murine tumor. Osaka City Med J, 1998. 44(1): p. 73-83.

55. Ishida, O., et al., Targeting chemotherapy to solid tumors with long-circulating thermosensitive liposomes and local hyperthermia. Jpn J Cancer Res, 2000.

91(1): p. 118-26.

56. Tiwari, S.B., R.M. Pai, and N. Udupa, Temperature sensitive liposomes of

plumbagin: characterization and in vivo evaluation in mice bearing melanoma B16F1. J Drug Target, 2002. 10(8): p. 585-91. 57. Aoki, H., et al., Therapeutic efficacy of targeting chemotherapy using local hyperthermia and thermosensitive liposome: evaluation of drug distribution in a rat glioma model. Int J Hyperthermia, 2004. 20(6): p. 595-605.

58. Dvorak, J., et al., Liposomal doxorubicin combined with regional hyperthermia: reducing systemic toxicity and improving locoregional efficacy in the treatment of solid tumors. J Chemother, 2004. 16 Suppl 5: p. 34-6.

59. Han, H.D., et al., Hyperthermia-induced antitumor activity of thermosensitive polymer modified temperature-sensitive liposomes. J Pharm Sci, 2006. 95(9): p. 1909-17.

60. Banno, B., et al., The functional roles of poly(ethylene glycol)-lipid and

lysolipid in the drug retention and release from lysolipid-containing thermosensitive liposomes in vitro and in vivo. J Pharm Sci. 99(5): p. 2295-308.

61. Mills, J.K. and D. Needham, Lysolipid incorporation in

dipalmitoylphosphatidylcholine bilayer membranes enhances the ion permeability and drug release rates at the membrane phase transition. Biochim Biophys Acta, 2005. 1716(2): p. 77-96. 62. de Smet, M., et al., Temperature-sensitive liposomes for doxorubicin delivery under MRI guidance. J Control Release. 143(1): p. 120-7.

63. Woo, J., et al., Use of a passive equilibration methodology to encapsulate

cisplatin into preformed thermosensitive liposomes. Int J Pharm, 2008. 349(1- 2): p. 38-46.

64. Ponce, A.M., et al., Hyperthermia mediated liposomal drug delivery. Int J

Hyperthermia, 2006. 22(3): p. 205-13.

65. Maeda, H., Tumor-selective delivery of macromolecular drugs via the EPR effect: background and future prospects. Bioconjug Chem. 21(5): p. 797-802. 66. Pardakhty, A., J. Varshosaz, and A. Rouholamini, In vitro study of

polyoxyethylene alkyl ether niosomes for delivery of insulin. Int J Pharm, 2007. 328(2): p. 130-41.

67. Liu, K., et al., Preparation and characterization of 10-hydroxycamptothecin loaded nanostructured lipid carriers. Drug Dev Ind Pharm, 2008. 34(5): p. 465- 71.

68. Pietkiewicz, J., et al., New approach to hydrophobic cyanine-type

photosensitizer delivery using polymeric oil-cored nanocarriers : hemolytic activity, in vitro cytotoxicity and localization in cancer cells. Eur J Pharm Sci, 2010. 39(5): p. 322-35. 69. Davis, S.S., et al., Use of adjuvants for enhancement of rectal absorption of cefoxitin in humans. Antimicrob Agents Chemother, 1985. 28(2): p. 211-5.

70. Bartlett, G.R., Colorimetric assay methods for free and phosphorylated glyceric acids. J Biol Chem, 1959. 234(3): p. 469-71.

71. Li, S.D. and S.B. Howell, CD44-targeted microparticles for delivery of

cisplatin to peritoneal metastases. Mol Pharm. 7(1): p. 280-90. 72. Lee, H., et al., Methoxy poly(ethylene glycol)-block-poly(delta-valerolactone) copolymer micelles for formulation of hydrophobic drugs. Biomacromolecules, 2005. 6(6): p. 3119-28.

73. Allen, T.M., P. Sapra, and E. Moase, Use of the post-insertion method for the formation of ligand-coupled liposomes. Cell Mol Biol Lett, 2002. 7(3): p. 889-

94.

74. MacKay, J.A., et al., Self-assembling chimeric polypeptide-doxorubicin

conjugate nanoparticles that abolish tumours after a single injection. Nat Mater, 2009. 8(12): p. 993-9. 75. Bradford, M.M., A rapid and sensitive method for the quantitation of

microgram quantities of protein utilizing the principle of protein-dye binding. Anal Biochem, 1976. 72: p. 248-54.

76. Szebeni, J., et al., The interaction of liposomes with the complement system: in vitro and in vivo assays. Methods Enzymol, 2003. 373: p. 136-54. 77. Goard, C.A., et al., Differential interactions between statins and P- glycoprotein: Implications for exploiting statins as anticancer agents. Int J Cancer.

78. Mclntyre, J.C. and R.G. Sleight, Fluorescence assay for phospholipid

membrane asymmetry. Biochemistry, 1991. 30(51): p. 11819-27. 79. Andresen, T.L., et al., Triggered activation and release of liposomal prodrugs and drugs in cancer tissue by secretory phospholipase A2. Curr Drug Deliv, 2005. 2(4): p. 353-62.

80. Gaber, M.H., et al., Thermosensitive sterically stabilized liposomes: formulation and in vitro studies on mechanism of doxorubicin release by bovine serum and human plasma. Pharm Res, 1995. 12(10): p. 1407-16. 81. Yuyama, Y., et al., Potential usage of thermosensitive liposomes for site- specific delivery of cytokines. Cancer Lett, 2000. 155(1): p. 71-7.

82. Chen, Q., et al., Tumor microvascular permeability is a key determinant for antivascular effects of doxorubicin encapsulated in a temperature sensitive liposome. Int J Hyperthermia, 2008. 24(6): p. 475-82.

83. Chen, Q., et al., Targeting tumor microvessels using doxorubicin encapsulated in a novel thermosensitive liposome. Mol Cancer Ther, 2004. 3(10): p. 1311-7.

84. Ponce, A.M., et al., Magnetic resonance imaging of temperature-sensitive liposome release: drug dose painting and antitumor effects. J Natl Cancer Inst, 2007. 99(1): p. 53-63.

85. Viglianti, B.L., et al., In vivo monitoring of tissue pharmacokinetics of

liposome/drug using MRI: illustration of targeted delivery. Magn Reson Med, 2004. 51(6): p. 1153-62.

86. Ickenstein, L.M., et al., Disc formation in cholesterol-free liposomes during phase transition. Biochim Biophys Acta, 2003. 1614(2): p. 135-8.

87. Sandstrom, M.C., et al., Effects of lipid segregation and lysolipid dissociation on drug release from thermosensitive liposomes. J Control Release, 2005. 107(1): p. 131-42.

88. Ickenstein, L.M., et al., Effects of phospholipid hydrolysis on the aggregate structure in DPPC/DSPE-PEG2000 liposome preparations after gel to liquid crystalline phase transition. Biochim Biophys Acta, 2006. 1758(2): p. 171-80.

89. Li, L., et al., Triggered content release from optimized stealth thermosensitive liposomes using mild hyperthermia. J Control Release. 143(2): p. 274-9.

90. Banno, B., et al., The functional roles of poly(ethylene glycol)-lipid and

lysolipid in the drug retention and release from lysolipid-containing thermosensitive liposomes in vitro and in vivo. J Pharm Sci, 2010. 99(5): p. 2295-308.

91. Brown, S.L., et al., Observations of thermal gradients in perfused tissues during water bath heating. Int J Hyperthermia, 1992. 8(2): p. 275-87. 92. Wust, P., et al., Implications of clinical RF hyperthermia on protection limits in the RF range. Health Phys, 2007. 92(6): p. 565-73.

93. Chopra, R., et al., An MRI-compatible system for focused ultrasound

experiments in small animal models. Med Phys, 2009. 36(5): p. 1867-74.

94. Floyd, J.D., et al., Cardiotoxicity of cancer therapy. J Clin Oncol, 2005. 23(30): p. 7685-96.

95. Peller, M., et al., MR characterization of mild hyperthermia-induced

gadodiamide release from thermosensitive liposomes in solid tumors. Invest Radiol, 2008. 43(12): p. 877-92.

96. Chiu, G.N., et al., Encapsulation of doxorubicin into thermosensitive liposomes via complexation with the transition metal manganese. J Control Release, 2005.

104(2): p. 271-88.

97. Abraham, S.A., et al., Formation of transition metal-doxorubicin complexes inside liposomes. Biochim Biophys Acta, 2002. 1565(1): p. 41-54.

98. Hasebroock, K.M. and N.J. Serkova, Toxicity ofMRIand CT contrast agents.

Expert Opin Drug Metab Toxicol, 2009. 5(4): p. 403-16.

99. Dong, X., et al., Development of new lipid-based paclitaxel nanoparticles using sequential simplex optimization. Eur J Pharm Biopharm, 2009. 72(1): p. 9-17.

Claims

CLAIMS:

1. A liposome comprising a lipid bilayer comprising l,2-dipalmitoyl-sn-glycero-3- phosphatidylcholine (DPPC) and a compound of formula:

C₁₇H₃5(CH₂)_p(CO)_q(OCH₂CH₂)_nOH

(I) wherein

p is an integer selected from 0 or 1;

q is an integer selected from 0 or 1;

p + q = 1 ; and

n is an integer selected from about 10 to about 100.

2. The liposome of claim 1 , wherein n is about 20.

3. The liposome of claim 1, wherein the lipid bilayer consists essentially of DPPC and Brij78, and preferably consists of DPPC and Brij78.

4. The liposome of any one of claims 1-3, wherein the diameter of the liposome is from about 30 ran to about 250 nm.

5. The liposome of any one of claims 1-4, further comprising at least one active agent entrapped in the interior space of the liposome, preferably a diagnostic agent or therapeutic agent.

6. The liposome of claim 5, wherein the active agent is hydrophilic and amphipathic.

7. The liposome of claim 5, wherein the diagnostic agent is gadolinium.

8. The liposome of claim 5, wherein the active agent is an anticancer agent.

9. The liposome of claim 8, wherein the active agent is selected from doxorubicin or gemcitabine.

10. The liposome of any one of claims 5-9, wherein the active agent is loaded into the liposome via a pH gradient.

11. The liposome of claim 10, wherein the molar ratio of DPPC to the compound of formula I is 99: 1 to 92:8.

12. The liposome of claim 11 , wherein the molar ratio of DPPC to the compound of formula I is at least 99: 1.

13. The liposome of claim 12, wherein the molar ratio of DPPC to the compound of formula I is at least 98:2.

14. The liposome of claim 13, wherein the molar ratio of DPPC to the compound of formula I is at least 96:4.

15. The liposome of claim 14, wherein the molar ratio of DPPC to the compound of formula I is about 96:4.

16. The liposome of claim 15, wherein the molar ratio of DPPC to the compound of formula I is at least 92:8.

17. The liposome of any one of claims 5-9, wherein the active agent is loaded into the liposome via a copper ion gradient.

18. The liposome of claim 17, wherein the molar ratio of DPPC to the compound of formula I is 84: 16 to 68:32.

19. The liposome of any one of claims 5-16, wherein the active agent is mixed with the liposome suspension at an active agent: lipid ratio of 1 :5 to 1 :40 w/w.

20. The liposome of claim 17 or 18, wherein the active agent is mixed with the liposome suspension at an active agent: lipid ratio of 1 :5.

21. The liposome of any one of claims 1 -20, having a T_m of about 42°C.

22. The liposome of any one of claims 5-21, wherein there is less than 20% leakage of the active agent at 37°C in 30 min, preferably less than 5% leakage of the active agent at 37°C.

23. The liposome of any one of claims 5-22, wherein there is substantially full release of the active agent in less than 10 min, preferably in less than 3 min, when the liposome is heated to 40°C .

24. The liposome of any one of claims 5-22, wherein there is substantially full release of the active agent in less than 10 min, preferably in less than 3 min, when the liposome is heated to 42°C.

25. The liposome of any one of claims 5-24, for delivery of the active agent to a target area in a patient.

26. A method of delivering an active agent to a target area in a patient, comprising: a. administering to the patient the liposome of any one of claims 5-25; b. heating the target area to at least 40°C for at least 3 min.

27. The method of claim 26, wherein the target area is heated to at least 42°C for at least 3 min.

28. Use of the liposome of any one of claims 5-25 for delivering the active agent to a target area in a patient, wherein the active agent is entrapped in the interior space of the liposome.

29. Use of the liposome of any one of claims 5-25 in the preparation of a medicament containing the active agent to be delivered to a target area in a patient.

30. A method for treating breast cancer in a patient in need thereof comprising administering to said patient a therapeutically effective amount of a liposome of claim 8 or 9.

31. A method for treating pancreatic cancer in a patient in need thereof comprising administering to said patient a therapeutically effective amount of a liposome of claim 8 or 9.

32. A method for treating lung cancer in a patient in need thereof comprising administering to said patient a therapeutically effective amount of a liposome of claim 8 or 9.

33. A method for treating multidrug resistant cancer in a patient in need thereof comprising administering to said patient a therapeutically effective amount of a liposome of claim 8 or 9.

34. The liposome of any one of claims 1-25, further comprising a targeting moiety.

35. The liposome of claim 34, wherein the targeting moiety is an RGD targeting sequence.

36. The liposome of claim 34 or 35, wherein the targeting moiety is covalently bound to the compound of formula I.

37. A liposome comprising a lipid bilayer comprising l,2-dipalmitoyl-5«-glycero- 3-phosphatidylcholine (DPPC) and Brij 56.

38. The liposome of claim 37, wherein the lipid bilayer consists essentially of DPPC and Brij 56, and preferably consists of DPPC and Brij 56.

39. A liposome comprising a lipid bilayer comprising l,2-dipalmitoyl-s«-glycero- 3-phosphatidylcholine (DPPC) and a compound selected from Brij 76, Brij700,

Myrj52, Myrj53, or Myrj59.

40. The liposome of claim 39, wherein the lipid bilayer consists essentially of DPPC and a compound selected from Brij 76, Brij^'700, Myrj52, Myrj53, and Myrj59, and preferably consists of DPPC and a compound selected from Brij 76, Brij700, Myrj52, Myrj53, or Myrj59. The liposome of any one of claims 37-40, further comprising at least one active agent entrapped in the interior space of the liposome, preferably a diagnostic agent or therapeutic agent.