MAGNETOENZYMATIC CARRIER SYSTEM FOR IMAGING AND TARGETED DELIVERY AND RELEASE OF ACTIVE AGENTS

Abstract

The present invention relates to a composition comprising liposomes, wherein the liposomes preferably comprise phospholipids, magnetic nanoparticles, an imaging agent and an imaging label, and at least one active agent. The liposomes are capable of releasing the at least one active agent into a target environment by the concomitant action of a phospholipase that is able to degrade at least one of said one or more phospholipids and an alternating magnetic field.

Claims

1. A composition comprising liposomes, the liposomes comprising: (a) phospholipids, the phospholipids comprising sphingomyelin (SM), (b) magnetic nanoparticles, and one or both of (c) an imaging agent and/or an imaging label and (d) at least one active agent, wherein the liposomes are capable of releasing an imaging agent and/or an active agent into a target environment by the concomitant action of sphingomyelinase (SMase) and an alternating magnetic field (AMF).

2. The composition of claim 1, wherein the sphingomyelin content of the liposomes is in the range of 10 mol % to 45 mol %.

3. The composition of claim 1, wherein the magnetic nanoparticles have a core diameter of between 2 nm and 10 nm.

4. The composition of claim 1, wherein the magnetic nanoparticles comprise a metal, metal alloy, metal oxide, metal hydroxide, mixed metal oxide, mixed metal hydroxide, metal nitride, metal sulfide, mixed metal nitride, or mixed metal sulfide, wherein the metal is selected from the group consisting of chrome, manganese, iron, cobalt and nickel.

5. The composition of claim 1, wherein the liposomes comprise at least one imaging agent selected from the group consisting of radioactive isotopes, contrast agents, fluorophores, and PET labels.

6. The composition of claim 1, wherein the active agent is a cytostatic agent.

7. The composition of claim 1, wherein the liposomes comprise at least one therapeutic agent, at least one imaging agent, and an imaging label.

8. A composition according to claim 1, for use in therapy and/or diagnostics.

9. The composition according to claim 8, wherein the use in therapy comprises the use of the composition for the targeted delivery and release of an active agent to a target site, and the use in diagnostics comprises the use of the composition in imaging.

10. A composition according to claim 1, for use in the treatment of cancer and/or for use in diagnosis of cancer.

11. A method of delivering and releasing an active agent to a target site, comprising the following steps: (a) administering a composition according to claim 1 to a target site, wherein the composition comprises at least one active agent; and (b) applying energy to liposomes in the composition and/or increasing SMase activity at the target site, thereby releasing the at least one active agent at the target site.

12. The method of claim 11, wherein the step of applying energy to the liposomes results in a phase transition and/or a change in the liposomal membrane structure and/or heating of the liposomes to effect release of the at least one active agent.

13. The method of claim 11, wherein in step (b) the energy is applied by applying an alternating magnetic field to the target site.

14. The method of claim 11, wherein the SMase activity is increased by administering exogenous SMase or by inducing endogenous SMase activity so as to effect release of the at least one active agent.

15. The method of claim 14, wherein the endogenous SMase activity is induced by irradiation, exposure to UV light, application of heat, or administration of chemotherapeutic agents or cell stress inducers.

Description

BRIEF DESCRIPTION OF THE DRAWINGS

[0047] FIG. 1 schematically shows a magnetoenzymatic SM liposome according to the present invention. (A) Schematic illustration of SM liposomes containing Fe.sub.3O.sub.4-nanoparticles embedded in the lipid membrane (dark grey globular structures within the lipid bilayer liposomal membrane) and a payload of imaging molecules (*) and/or drug molecules () inside the liposome. (B) TEM image of SM liposome (without membrane staining) with 5 nm Fe.sub.3O.sub.4 nanoparticles agglomerated in the lipid membrane. C. TEM image of SM liposome without Fe.sub.3O.sub.4 nanoparticles. The liposomal membrane was stained with uranyl acetate.

[0048] FIG. 2 shows validation of encapsulation of (treatment and/or imaging) agents and assessment of liposome size. Data for SM liposomes incorporating different types of anticancer drugs: (A) 0.5 mg/ml doxorubicin, (B) 4 mg/ml cisplatin, (C) 0.7 mg/ml paclitaxel; or for SM liposomes incorporating different types of imaging agents: (D) 0.5 mg/ml indocyanine green (ICG). Upper panels: typical elution curve from a PD-10 column for DSPC:SM:DOTAP:chol (20:30:20:30) liposomes with Fe.sub.3O.sub.4 nanoparticles (coated with catechol-PEG400 and carboxy group) on the membrane, containing the respective payload; fraction number three at volume 4.2 ml, containing the intact liposomes, was collected for downstream use. Lower panels: distribution of the mean radius of the SM liposomes as measured by dynamic light scattering analysis.

[0049] FIG. 3 shows the assessment of fluorescent signals of free and encapsulated agents. (A) Effect of breaking of the lipid membrane for doxorubicine in SM liposomes. Upper horizontal and vertical Eppendorf reaction vessels: SM Liposomes with doxorubicine and 5 nm Fe.sub.3O.sub.4 nanoparticles before (horizontal Eppendorf reaction vessel) and after (vertical Eppendorf reaction vessel) SM liposomes had been broken up with a 49% EtOH solution. Lower horizontal and vertical Eppendorf reaction vessels: SM liposomes with doxorubicine (without Fe.sub.3O.sub.4 nanoparticles) before (horizontal Eppendorf reaction vessel) and after (vertical Eppendorf reaction vessel) SM liposomes had been broken up with a 49% EtOH solution. (B) Effect of the breaking of the lipid membrane for ICG (with 5 nm Fe.sub.3O.sub.4 nanoparticles) in SM liposomes before (bottom right) and after (top left) treatment with 49% EtOH solution to break the lipid membrane.

[0050] FIG. 4 shows the assessment of permeation of a SM liposome by SMase. (A) Liposome disruption in vitro by SMase. Lipid content 20 mol, SM content 6 mol. The relative fluorescence intensity (RFI) detected by FRET DNA hairpin assay is shown as a function of different levels of SMase exposure. (B) Liposome disruption in vitro by aSMase secretion by Human Aortic Endothelial Cells (HAoEC) induced by radiation stress with exposure levels ranging from 0 to 15 Gy. aSMase activity was measured using the sphingomyelinase activity assay kit from Cayman Chemical.

[0051] FIG. 5 shows the specific power absorption (SPA, W/g) of Fe.sub.3O.sub.4 nanoparticles in a liposome environment and free in water and exposed to an AMF of 23 mT at 828 kHz. Various concentrations of 5 nm Fe.sub.3O.sub.4 nanoparticles embedded in ICG loaded SM liposomes (1.2 mg/ml, 0.6 mg/ml, 0.3 mg/ml, and 0.15 mg/ml) were compared with 20 nm and 5 nm Fe.sub.3O.sub.4 nanoparticles free in water (the two leftmost columns). At optimum concentrations of 0.3 mg/ml, the embedded 5 nm nanoparticles demonstrate 10 times higher SPA values compared to 5 nm particles free in water. Upon addition of SMase in amounts sufficient to destabilize the membrane of the liposomes and release the iron oxides from the membranes, the SPA of the 5 nm particles was reduced below the detection limit.

[0052] FIG. 6 shows the results obtained for SM liposomes in a leakage assay. The two bars on the left show the results for SM liposomes containing Fe.sub.3O.sub.4-nanoparticles (DSPC:SM:DOTAP:chol (20:30:20:30) mol/mol) with AMF exposure (1.5 mT, 100 kHz), and the two bars on the right show the result for SM liposomes without Fe.sub.3O.sub.4-nanoparticles (SM/DSPC/chol (30:40:30) mol/mol) and without AMF exposure. The white bars indicate no SMase exposure, and the dark bars indicate 0.8 U/ml SMase exposure. Exposure to both SMase and AFM resulted in a leakage rate that was about three times higher compared to only SMase or only AMF exposure.

[0053] FIG. 7 shows the results obtained for different drug formulations in a liposome leakage assay. The SM liposomes were made of DSPC:SM:DOTAP:chol (20:30:20:30 mol/mol) with 5 nm Fe.sub.3O.sub.4 nanoparticles (0.6 mg/ml). (A) Doxorubicine (0.5 mg/ml) formulation. Left bar: without AMF and SMase exposure, middle bar: AMF only (5 mT, 229 kHz), right bar: AMF (5 mT, 229 kHz) after 30 min SMase (0.8 U/ml) pretreatment. (B) Cisplatin (4 mg/ml) formulation. Left bar: without AMF and SMase exposure, middle bar: AMF only (23 mT, 823 kHz), right bar: AMF (23 mT, 823 kHz) after 30 min SMase (0.8 U/ml) pretreatment. (C) Paclitaxel (0.7 mg/ml) formulation. Left bar: without AMF and SMase exposure, right bar: AMF (23 mT, 823 kHz) after 30 min SMase (0.8 U/ml) pre-treatment. (D) Leakage of doxorubicine (0.5 mg/ml) from SM liposomes after 1 h incubation with PC-3 cells exposed to aSMase induced by 36 Gy radiation (left) and without irradiation (right).

[0054] FIG. 8 illustrates the treatment of cancer in a mouse model using SM liposomes according to the present invention. (A) Schematic drawing of the experimental set up used in AMF experiments for the treatment of oral squamous cell carcinoma (SCC). The top of the coil is placed near the tumor and the tumor is exposed to AMF treatment. (B) Visualization of the release of an imaging marker from SM liposomes in an orthotopic mouse model of SCC4 oral squamous cell carcinoma. ICG encapsulated in 5 nm Fe.sub.3O.sub.4 nanoparticles containing SM liposomes does not yield a visible fluorescent signal prior to AMF treatment (left), but leads to a clearly visible fluorescent signal after release due to the AMF treatment (20 min, 1.5 mT, 100 kHz).

DETAILED DESCRIPTION OF THE INVENTION

[0055] The present invention combines enzymatic action together with AMF triggered magnetic nanoparticle containing liposome release, resulting in a more specific and more selective release than can be achieved at physiological SMase concentrations found in various diseases and with AMF ranges suitable for use in clinical settings.

[0056] The liposome carriers of the invention feature a specific externally controllable release mechanism which allows the rupture of the liposome when being in correct place and when needed, leading to release of the payload. For this purpose two different mechanisms are combined: (a) degradation of sphingomyelin as a lipid constituent of the liposome by enzymatic action of SMase secreted by pathologic cells, e.g. in cancer and in inflammation, and (b) magnetic nanoparticles used for destabilizing the liposome membrane by the action of AMF, thereby inducing leakiness.

[0057] Furthermore, acid or neutral SMase secretion can be enhanced by localized irradiation, various stress stimulators like lipopolysaccharides, disruption of integrin signaling, UV-light, heat, oxidative stress and accumulation of Cu.sup.2+, participation of platelet-activating factor, chemotherapeutic agents (e.g., cisplatin, etoposide, gemcitabine, fenretinide, paclitaxel, rituximab, daunorubicin, Ara-C, and doxorubicin), or it can be activated via TNF-receptor superfamily members (Fas, CD40, DRS, and TNF). These stress stimulators can be used to induce the increase of SMase at the site of illness which may release further compounds, which can again activate more aSMase and thus multiply the drug effect at the site of disease.

[0058] Within the context of the present invention, SMase from external sources may also be used. There are several parasites, bacterial strains and viruses that utilize SMase in their life cycle. This innovation allows the usage of these non-human sources of SMase to be detected and utilized for targeted release and imaging. This includes for example SMase secreting bacteria like B. cereus (Oda et., 2010), C. pneumonia (Penate Medina et al., 2014), and S. aureus (Hedstrm and Malmqvist, 1982), pathogens observed frequently in hospital environments, as contaminants in food and drinks and on material surfaces including, e.g., implants.

[0059] For the sphingomyelin (SM) liposomes according to the invention, commonly occurring SMs such as 16:0 SM, 16:0 SM (d18:1116:0), 17:0 SM (d18:1117:0), 18:0 SM (d18:1/18:0) may be used. The SM content of the liposomes should be in the range of 10 to 45 mol %. However, the upper limit of the SM content may also be as high as 50 to 60 mol %. In the context of the present invention, the SM content is preferably in the range of 10 to 45 mol % or 15 to 40 mol %, e.g., 20 to 35 mol % or 25 to 35 mol %. The amount of SM as substrate for SMase has to be sufficient to induce creation of ceramide rich platforms that make the liposomes sensitive to the AMF treatment. A level of about 5% ceramide in the membrane has to be reached for ceramide-rich domains to form (Sot et al., 2006) and 10% of ceramide in the membrane for vesicle budding to happen (Nurminen et al., 2002).

[0060] Furthermore, other liposome forming lipids may be used such as those selected from the group consisting of phospholipids, tocopherols, sterols (e.g., cholesterol and derivatives thereof, ergosterol and derivatives thereof, lanosterol and derivatives thereof), glycoproteins, and mixtures thereof. The phospholipids may be selected from the group consisting of: (1) phosphatidylcholine (PC) (e.g., dioleoylphosphatidylcholine (DOPC), dilinoleoylphosphatidylcholine (DLPC), dimyristoylphosphatidylcholine (DMPC), dipalmitoylphophatidylcholine (DPPC), disaturated-phosphatidylcholine (DSPC), dioleoylphosphatidylcholine (DOPC), egg PC, hydrogenated soy PC); (2) phosphatidylglycerol (PG) (e.g., dimyristoylphosphatidylglycerol (DMPG), dipalmitoylphosphatidylglycerol (DPPG), distearoylphosphatidylglycerol (DSPG), egg PG); (3) phosphatidylethanolamine (e.g., dimethylphosphatidylethanolamine (DMPG), dipalmitoylphosphatidylethanolamine (DPPE), distearoylphosphatidylethanolamine (DSPG), dioleoylphosphatidyl ethanolamine (DOPE)); (4) phosphatidylserine (PS) (e.g., dioleoylphosphatidylserine (DOPS)); (5) phospatidic acid (e.g., dimyristoylphosphatidic acid (DMPA), dipalmitoylphosphatidic acid (DPPA), distearoylphosphatidic acid (DSPA)); (6) PEGylated derivatives of (1) to (5); (6) other lipids (e.g., N-(2,3-di-(9-(Z)-octadecenyloxy)-prop-1-yl-N,N,N-trimethylammonium chloride (DOTMA) and 1,2-bis(oleoyloxy)-3-(trimethylammonio)propane (DOTAP)), and combinations thereof.

[0061] Suitable liposomes for use herein may in particular comprise at least one sphingomyelin in an amount of 10 to 60 mol % (e.g., 10 to 50 mol %, 10 to 45 mol %, 15 to 40 mol %, 20 to 35 mol % or 25 to 35 mol %) and at least one lipid selected from the group consisting of: phosphatidylcholine (PC) (e.g., DPPC, DSPC, DOPC and DMPC) in a total amount of about 10 to 85 mol % (e.g., 15 to 50 mol %, 15 to 40 mol % or 20 to 35 mol %), cholesterol (or derivatives thereof) in an amount of 5 to 40 mol % (e.g., 10 to 35 mol % or 20 to 30 mol %), phospatidylethanolamine (e.g., DMPE, DPPE, DSPE) in a total amount of 0.1 to 30 mol % (e.g., 0.1 to 20 mol %, or 1 to 10 mol %), phosphatidylserine (e.g., DOPS) in a total amount of 1 to 35% (e.g., 5 to 25% or 10 to 20 mol %), DOTAP and/or DOTMA in a total amount of 5 to 40 mol % (e.g., 5 to 30 mol %, 5 to 20 mol % or 10 to 25 mol %), and combinations thereof.

[0062] Preferably, the other liposome forming lipids (i.e. lipids besides the SM lipids as defined above) may be: cholesterol 10-30% mol %, phosphatidycholine (PC) 25-80 mol %, cationic lipids (DOTAP) 5-20 mol %, phosphatidylethanolamine (PE) 0.1-10 mol %, or phosphatidylserine (PS) 5-20 mol %. The preferred size of the liposomes is from 50 to 300 nm, more preferably from 80 to 250 nm, which allows sterile filtration prior to in vivo injection (using standard 0.22 micron sterile filters), but it may range up to 800 nm.

[0063] Accordingly, in the course of action of SMase on mixed SM/phosphatidylcholine membranes the generated ceramide forms microdomains (Nurminen et al., 2002). In terms of physical chemistry this represents an isothermal transition triggered by an enzymatic reaction, causing a shift in the lipid phase diagram from the fluid disordered phase into the two phase region, consisting of the fluid disordered and the solid ordered phases, the latter being enriched in ceramide. This isothermal transition will permanently alter the state of the liposome before and after enzymatic cleavage. Before SMase action the liposome is rigid and non-leaky, after SMase action the liposome is fragile and unstable. This instability of the lipid membrane allows to increase leakage by applying heat or by inducing increasing lateral movement on the membrane resulting in further destabilization of the membrane. These effects can be generated by AMF induction through beads agglomerated in the membrane and/or by external heat sources like ultrasound, laser or heating probes.

[0064] The liposomes, also referred to as primary carrier, according to the invention may carry medical agents, diagnostic agents, nutritional agent, a radiation sensitizer, a contrast agent, an enzyme, nucleic acid, an antibody, a growth factor, a protein, a peptide, carbohydrate, a targeting group or combinations of those.

[0065] For the magnetoenzymatically controlled active agent release by an AMF upon SMase action, the core size of the Fe.sub.3O.sub.4 nanoparticles may be in the range of 2 to 10 nm, preferably 5 nm. Also other magnetic, superparamagnetic, paramagnetic and ferromagnetic particles in the same size range can be used. Small nanoparticles incorporate spontaneously in the membrane. Carboxy coated 5 nm Fe.sub.3O.sub.4 nanoparticles associate spontaneously to the membrane (5 nm particles with cathechol-PEG-COOH group). When nanoparticles in the bloodstream have a size of 2 to 10 nm they can be cleared through the kidney. The iron oxide nanoparticles are therefore preferably 10 nm or smaller, but larger or equal to 2 nm. Kidney clearance is difficult or nearly impossible if the particle size is larger than 10 nm in mice. In humans the exact shape and size of the kidney threshold has not been established but it is thought to be comparable and vary with the age and health of the patient.

[0066] The magnetic field strength for drug release by AMF may be in the range of 0.3 mT to 60 mT, and the field modulation frequency in the range of 100 to 1000 kHz. The application of the AMF may be totally local and can be done after or during imaging, enabling total control of the process. On the other hand, the fact that SMase activity is only associated with severe disease conditions provides an insurance against agent release in false-positive areas of disease or off-target accumulation that will naturally arise from the uneven distribution of the liposomes in the body.

[0067] The enhanced sensitivity to AMF in the presence of SMase brings advantages for clinical application since it may permit the use of surface coils for generating the AMF. Small portable coils allow selective AMF induction at the sites of the tumor only. This spares the other parts of the body from drug release from the liposomes and from heating. It is known that energy dissipation, i.e. heat generation, is linearly dependent on the frequency, it increases with the second power of the magnetic field, and for surface coils it decreases with the third power of the distance. While this allows for more spatially selective exposure it may also limit the usage of small surface coils outside the body of the patients. Due to the synergistic effect of SMase and AMF the effective magnetic field used can be lower and this way the usability of the surface coils for initiating liposomal drug release is enhanced. While it also would be possible to place a patient inside a magnetic coil this requires higher magnetic fields which, due to medical risks, might be increase procedural complexity and may be contraindicated for some patients (e.g. with pacemakers) and medical personnel.

[0068] Unlike magnetoenzyme sensitive SM liposomes thatwhen exposed to AMFwill only open in the vicinity of pathologies expressing SMase, thermosensitive liposomes described in the state-of-the-art do not demonstrate such a selective releaseall the liposomes in the magnetic field will become permeable once a certain level of energy is disposed. Moreover, in those systems the heating of the surrounding tissues continues even after liposomes have been broken as long as AMF exposure continues. Furthermore, thermosensitive liposomes, if equipped with imaging markers that change emission patterns upon release of the agent, will yield signals independent of the SMase state of the diseased target. Therefore, they lack the site specific SMase mediated information about disease activity reflected in the signal of the imaging marker released, a critical advantage in the assessment of disease status allowing to regulate and thus optimize the aggressiveness of therapy.

[0069] The combination of image control and disease mediated release rates provides the components required to realize a feedback loop system for adaptive and controlled therapy. Since AMF exposure allows to adjust release sensitivity to detectable levels of image signal modulation for varying levels of SMase expression this invention can also be used to follow the pathophysiological status of the cancer or the severity of inflammation by monitoring the release from SM liposomes caused by secreted SMase in the site of disease. The level of AMF required to detect a certain level of release of a marker would reflect SMase expression levels, e.g. associated with apoptosis, and thus can indicate both the effectiveness of the drug treatment and/or the severity of the illness.

[0070] This invention can also bring about improved methods for clearing the drug that did not reach the intended target pathology. For most cancer drugs 95.0 to 99.9% of the injected drug never associate to the target tissue and target receptors (for an example involving doxorubicine see Gasselhuber et al., 2012). This off-target drug load causes adverse effects that limit the usability of the drug. The invention enables the encapsulated drug to circulate in the blood without interacting with off-target tissue. Clearance of off-target drug can be achieved in several different ways, e.g. by extracting the non-opened off-target iron oxide SM liposomes with a magnet from the blood (e.g. in a dialysis like setting).

[0071] The incorporation of iron nanoparticles in the liposomes was realized with the goal to improve the drug release efficacy and the better control of the drug release. An ideal nanoparticle would spontaneously associate to the membrane but dissociate after the liposome is made leaky because this would result in a self-limiting heating pattern, avoiding unnecessary heat generation that might lead to undesirable side effects. As nanoparticles for the magnetic action, the inventors used carboxyl functionalized Fe.sub.3O.sub.4 particles (catechol-PEG-COOH capped Fe.sub.3O.sub.4 particles, AC Diagnostics, Fayetteville, Ark., USA). Small iron nanoparticles incorporated spontaneously to the lipid films. This may be due the opposite charges of the nanoparticles and the DOPA in the liposomes. On the other hand, phosphate that can be found in the phospholipid head groups is known to work as a stabilizer of the iron nanoparticles. It is also known that catechols have a tendency to leave the nanoparticles and this way allow the phosphate ions to interact with the iron core.

[0072] A schematic illustration of SM liposomes containing iron nanoparticles embedded in the lipid membrane stabilized by the phosphates of the lipids and containing payload (FIG. 1A) shows that the payload active agent may be in the aqueous compartment inside the liposome. However, alternatively it also may be embedded in or attached to the membrane. Because isolated iron containing particles are too small to generate enough heat in aqueous solutions under AMF (Goya et al., 2008), they have to be clustered as depicted in FIG. 1B. This figure presents the TEM image of SM liposomes with 5 nm Fe.sub.3O.sub.4 nanoparticles. The agglomeration of the nanoparticles can be clearly seen. TEM imaging shows both clear association of the 5 nm nanoparticles on the membrane and large agglomerates in the membrane. The agglomerates also make the membrane to bulge. In the membrane of stained control liposomes without iron particles, smooth symmetrical contours are seen on the TEM image (FIG. 1C).

[0073] The addition of Fe.sub.3O.sub.4 nanoparticles does not affect leakage. In the cationic liposomes, the Fe.sub.3O.sub.4 nanoparticles associate on the liposomes even though they are stabilized by catechol-PEG 400. This might be due to the dissociation of the catechol PEG coating from Fe.sub.3O.sub.4 nanoparticles and association of the Fe.sub.3O.sub.4 nanoparticles with phosphate groups of the phospho- and sphingolipids. Phosphate is a known stabilizer of iron nanoparticles and in this sense the iron particles could be stabilized on the liposomal water lipid interface where the charged head groups are available without additional stabilizers like catechols. Binding of iron nanoparticles to the membrane might also be strengthened by the opposite surface charges of the liposome membrane and iron-nanoparticle surface. Association of iron nanoparticles on the liposomal membrane is dependent of the state of the membrane and is reversible. Particles can dissociate from the membrane if the state of the membrane changes, for example when ceramide rich micro-domains are formed in the membrane.

[0074] PEG 400 may be used as solubilizing linker to the stabilizer catechol with a total MW of 760. This stabilizer linker complex is one order of magnitude smaller than the known efficient steric stabilizers. It is also known that catechols are more unstable than nitrocatechols and can dissociate from nanoparticles. These two facts allow the particle to interact with the surrounding phosphates from the phospholipids. The nanoparticles that have catechol DOPA and PEGs shorter that 1000 Da have a tendency to loosely associate with phospholipids when DOPA PEG can be depleted and this may help the nanoparticles to agglomerate at the membrane (Isa et al., 2010, and Goldmann et al., 2010). It is possible to consider the use of other stabilizers like oleic acid, silane, phosphate, and dopamine, since these stabilizers may be modified with a great variety of groups, as long as the particle associates and dissociates to/from the membrane.

[0075] Typically, the more stable dispersants such as nitro-DOPA-PEG5000 provide a dense, thin layer which is sufficiently thick to prevent nanoparticle agglomeration. PEGs from 5000 Da to 10 000 Da make the most stable nanoparticles that do not allow agglomeration and interaction of the iron nanoparticle with the lipid film. Reimhult et al. (2012) used Fe-particles of the roughly the same size as in the present invention but they used magnetic nanoparticles, which are stabilized using nitrocatechols with high-affinity palmitoyl-anchors to establish a thin but very dense hydrophobic coating, which can be incorporated into a membrane. Nitrocatechols are more stable and form less radicals than catechols (Amstad et al., 2009, and Amstad et al., 2010). However, this nitrocatechol bonding and anchoring is extremely stable and almost irreversible and makes the nanoparticle elimination from the body extremely difficult because the palmitoyl anchors can anchor the complex to any membrane.

[0076] Commonly occurring SMs were used, which can be in a bilayer membrane of liposomes such as 16:0 SM, 16:0 SM (d18:1116:0), 17:0 SM (d18:1117:0), 18:0 SM (d18:1118:0). In the present invention the lipids used have been widely tested in clinical settings and thus are acceptable for clinical translation.

[0077] The SM liposomes containing drugs, ICG and iron nanoparticles were stable and the encapsulation efficiency was generally good. Doxorubicine (0.5 mg/ml) was encapsulated using a pH gradient and the encapsulation efficiency was high, as could be seen from the gel filtration elution curves. In a typical elution curve from a PD-10 column for DSPC:SM:DOTAP:chol (20:30:20:30) liposomes with Fe.sub.3O.sub.4 nanoparticles (coated with catechol-PEG400 and carboxy group), fraction number three contained the intact liposomes and was collected for further use (FIG. 2A). Cisplatin (4 mg/ml) encapsulation was performed by combining the drug with the lipids when the lipids and cisplatin were hydrated. Fraction number three contained the intact liposomes and was collected for further use (FIG. 2B) Paclitaxel (0.7 mg/ml) was encapsulated by dissolving paclitaxel in the lipid mix when in chloroform. Fraction number three contained the intact liposomes and was collected for further use (FIG. 2C). ICG (0.5 mg/ml) was encapsulated with a similar strategy and added after the formation of the lipid film. Fraction number three contained the intact liposomes and was collected for further use (FIG. 2D). Homogeneity and size distribution were checked by using dynamic light scattering (DLS).

[0078] To make the liposomes easier to break they can be created more fragile, more fluid like, by using lower amounts of cholesterol, and/or by using non saturated fatty acids in the lipid mixture. When these kinds of liposomes are used perhaps even smaller amounts of SMase, like found in infections, could be enough to break the liposomes but weaker liposomes have shorter storage half-life and are leaky in the blood stream. SM DSPC liposomes, on the other hand have large drug loading capacity, long shelf-life and robust circulation times. SM DSPC liposomes thus can be used to keep the liposomes always intact in physiological situations. Incorporation of iron nanoparticles makes them vulnerable only in case of exposure to magnetic fields that induce a phase transition. In this case, the initial sensitivity to low SMase concentrations may be reduced but better control of the release could be achieved.

[0079] Generally, the liposomes of or used in the present invention are not thermosensitive. The term not thermosensitive, as used herein, preferably means that the liposomes are not sensitive (i.e. are stable, in particular in the sense that no leakage of, i.e., drugs or other active agents, occurs) in the typical range of body temperatures (e.g., 36-38 C. or 36.5-37.5 C.).

[0080] Liposomes that are composed of lipids that have T.sub.m of 55 C., DSPC of 41 C. sphingomyelin and of 4 C. cholesterol DOTAP are rigid and non-thermosensitive. They are also stable, with a shelf-life of several weeks. If not exposed to SMase (e.g., aSMase), phase transitions of liposomes with, e.g., sphingomyelin, cholesterol, phosphatidylcholine and DOTAP occur between 60-75 C., depending on the SM and helper lipids (i.e. all lipids except SM) concentrations. These such liposomes are not thermosensitive. Hence in the absence of SMase exposure (with or without AMF), no drug release is possible. For iron oxide containing variants, the liposomes of this rigidity do not leak either, even under AMF exposure. After SMase treatment, heat induced either directly or via ultrasound, or other thermal device would work as well as AMF, but applying heat to inner parts of the body in a controlled selective manner is challenging. The SM liposomes described here permit permeabilization even at low energy AMF (i.e. energies that do not increase the tissue temperature over 40 C.). Thus this avoids heat and other damage (e.g. by exposure to higher levels of ionizing radiation to induce higher levels of SMase).

[0081] The level and spectral distribution of ICG fluorescence is dependent on ICG concentration and the local environment. In vitro, ICG fluorescence is brightest at a concentration of 80 g/ml and above that the fluorescence decreases rapidly when concentration increases (Mordon et al., 1998). In the experiments presented here the liposomal ICG concentration was 750 g/ml, but the actual local concentration was several fold higher because the ICG was trapped in liposomes and possibly in the liposomal bilayer. This leads to an efficient quenching in cationic liposomes (Hua et al., 2012). The same quenching phenomenon applies to the doxorubicine liposomes: the doxorubicine fluorescence increases 10-fold when released from liposomes. Doxorubicine fluorescence can be used to monitor the release (Kheirolomoom et al., 2010).

[0082] To test if the fluorescence properties of ICG and doxorubicine could be used to monitor whether the agents are free or encapsulated in liposomes the following experiment was carried out. Doxorubicine (with or without iron oxide nanoparticles) or ICG were encapsulated in liposomes. After the liposome membrane was broken by 49% ethanol, an increase in the fluorescence signal was observed for both the free doxorubicine (FIG. 3A) and the free ICG (FIG. 3B). This effect can be explained by the p-p orbital stacking effect (Hua et al., 2012). The effect thus indeed can be exploited to monitor the release from liposomes and also to detect pathologic sites of high SMase expression and AMF exposure is to be used to bring imaging signals to levels detectable and monitorable. Such monitoring can be very informative when performed during surgery: intra operative imaging using near infrared markers such as ICG encapsulated in SM liposomes may be used to detect tumor tissue in order to verify successful surgery, i.e. complete removal of all tumor tissue. These optical signals can also be helpful as input for steering robot surgery devices.

[0083] In order to show that SMase treatment can release the molecules entrapped in the lumen of liposomes a liposomal hairpin test was carried out. A dose dependent increase of the fluorescent signal associated with SMase exposure was observed when hairpin DNA was liberated from liposomes. With this liposomal hairpin model it could be shown that there is leakage from SM containing liposomes when SMase concentration reaches 0.5 U/ml (FIG. 4A). FIG. 4B demonstrates that the level of aSMase required for permeation of SM liposomes can be produced by cells, in this case human aortic endothelial cells (HAoEC), exposed to levels of ionizing radiation as applied in clinical conditions like in cancer radiation therapy.

[0084] The effect of the applied AMF on the iron containing nanoparticles in liposomal or water environment and with different nanoparticle concentrations and sizes was studied in more detail by analyzing specific power absorption (SPA) values (W/g). Also the effect on SPA of exposure of the iron containing liposome system to SMase was studied. ICG containing SM liposomes with various 5 nm iron nanoparticle concentrations were investigated. For comparison, 20 nm and 5 nm iron nanoparticles were analyzed free in water (FIG. 5). Three surprising observations were made.

[0085] First, a tenfold increase in the SPA values was measured when 5 nm particles were coupled to liposomes at optimal concentration levels (lipo 0.3 mg/ml, FIG. 5) vs. when they were free in water environment (5 nm Water, FIG. 5). When small 5 nm iron nanoparticles agglomerate in the lipid membrane they allow heating of the liposomes. In an optimum agglomerated state in the liposomal membrane 5 nm iron particles reach SPA values of about a third of the SPA values of 20 nm iron particles.

[0086] Second, the increase of the power absorption is dependent on the nanoparticle concentration in the liposomal environment. Still there is a limit: in a lipid membrane there is only space for limited amount of the agglomerated particles and this defines the upper limit for the SPA values. If more particles are added to the membrane they just are not optimally distributed to the membrane and that results in lower SPA values.

[0087] Third, the SPA effect for 5 nm iron containing nanoparticles in the liposomal membrane vanishes after exposure to SMase present in the solution (lipo 0.6 mg/ml SMase compared to lipo 0.6 mg/ml in FIG. 5). This could be explained by the formation of ceramide rich microdomains that can destroy the agglomeration of the particles within the membrane. Particles released from the bilayer to the surrounding aqueous environment would lose the supernormal heating capability brought about by agglomeration.

[0088] It was also studied if there is synergy for drug release when using both AMF and SMase. Relatively low magnetic fields were used and enzyme amounts were chosen to simulate physiological conditions. The leakage of the ICG from ICG containing liposome with and without AMF and SMase treatment was measured. While enzyme treatment alone or the AMF exposure could induce only a limited leakage effect, an unexpected synergistic 3-fold increase of leakage of ICG from liposomes was induced by combined enzyme and magnetic field exposure (FIG. 6).

[0089] Taken together, these various surprising observations reveal a very favorable drug release activation pattern for SM liposomes with nanoparticles of very small size (5 nm) that can agglomerate appropriately in the liposome membrane: a synergistic magnetoenzymatic effect permits reaching energy absorption levels high enough to cause leakage of the SM liposomes provided these are exposed to SMase. And once the liposomes have become leakier due to SMase exposure, the energy absorption is minimized, reducing any unnecessary further heat transfer to surrounding tissues after the payload has been released. This feature can be exploited to enhance the therapeutic window of AMF by either reducing potential side effects of heating or by permitting safe enhancement of energy absorption (adapting frequencies and magnetic field strength) for enhanced release of active drugs.

[0090] In order to confirm that release is not only feasible for imaging markers but also that the magnetoenzymatic carrier system can be used for drug release as well, additional experiments were performed. Doxorubicin, cisplatin, and paclitaxel were encapsulated and their release from the encapsulating SM liposomes by using the synergistic effect of combined SMase and AMF exposure was documented (FIG. 7). The experimental setup was similar as for the liposomal leakage assays described above, exposing the carrier to SMase and/or AMF treatment at different magnetic field strengths and frequencies.

[0091] FIG. 7A depicts results for the doxorubicine liposomal formulation for an AMF treatment at 5 mT with 229 kHz, while FIG. 7B shows similar results for the cisplatin liposomal formulation with an AMF treatment at 23 mT with 823 kHz. Additional exposure to 0.8 U/ml of SMase led to an increase in the release rate, documenting that relatively low magnetic field strength may be sufficient, which may facilitate clinical application.

[0092] In FIG. 7C results for yet another liposomal formulation are presented, in this case for paclitaxel with an AMF treatment at 23 mT with 823 kHz. In the release experiments of FIGS. 7A-C it has been shown that the drugs can be released by combined SMase and AMF exposure in a manner similar to the ICG release assays. The results vary for different drugs and are most impressive for the doxorubicine formulation, which showed a more than 10-fold increase in the case of additional SMase exposure, achieved at a relatively low magnetic field strength of 5 mT.

[0093] In the case of doxorubicine the leakage rate of liposomes when exposed to aSMase expressed by PC-3 cells by subjecting the cells to 36 Gy of ionizing radiation was investigated (FIG. 7D). Extracellular expression of SMase and SMase induced ceramide are important reactions to cell stress and apoptosis, commonly observed in inflammation and cancer. Therefore the finding that SMase can be used to break specific types of liposomes, i.e. SM liposomes, makes it an interesting target to design drug delivery systems for endogenously triggered drug release. 2 U/ml SMase are enough to break SM liposomes but it this is a level that is rarely reached in typical pathophysiological settings. It has been reported that SMase activity of 0.5 U/ml can lead to 80% cell death in colon HT-29 cancer cells (Colell et al., 2002).

[0094] FIG. 6 shows that 30 min incubation with 0.8 U/ml SMase is not enough to release the content alone, but the additional exposure to AMF leads to a leakage rate of more than 50%. This supports the use of SM liposomes loaded with iron containing nanoparticles as drug delivery vehicles: once accumulated at sites of pathologic expression of SMase, AMF treatment will lead to the release of their payload, including drugs and/or imaging markers.

[0095] Using an experimental setup as depicted in FIG. 8A it was tested whether the unquenching of ICG associated with the release from the liposome and observed in vitro (FIG. 3) could be observed in vivo. ICG loaded Fe.sub.3O.sub.4 nanoparticle containing SM liposomes were injected into the tail vein of SCC4 orthotopic xenograft bearing mouse. The liposomes circulated for 15 min, then the animal was exposed to AMF (20 min at 1.5 mT, 100 kHz). AMF treatment was directed to the tumor under the jaw, as depicted in FIG. 8A. Near infrared imaging was performed in a NightOWL fluorescence imaging chamber 1 hour after the injection. As control, mice were injected with the same liposomal composition but not subjected to AMF treatment. The leaked ICG could be clearly seen in the tumor area under the chin of the AMF treated mouse (FIG. 8B right) but not on the control mouse (FIG. 8B left). This phenomenon with ICG quenching when encapsulated in liposomes is shown in FIG. 3B.

[0096] The present invention will now be further illustrated by the following, non-limiting examples.

Example 1

Liposome Formulations

[0097] Lipids were purchased from Avanti Polar Lipids and SMase (from B. cereus) from Sigma Aldrich.

[0098] a) Doxorubicin Loaded Liposomes:

Liposomes were made using a lipid mixture consisting of a total of 20 M lipids DPPC/Cholesterol/SM/DOTAP or DSPC/cholesterol/SM/DOTAP (20:30:30:20 mol %, respectively). In order to encapsulate doxorubicin into the liposomes a pH-gradient was used. Lipids stored in chloroform were pipetted to a round bottomed flask, dried under nitrogen and lyophilized for 24 h to remove trace amounts of chloroform. Lipids were allowed to hydrate for 30 min, at 50 C. in the case of DPPC and 60 C. for DSPC, in 1 ml of 250 mM (NH.sub.4).sub.2SO.sub.4, 150 mM NaCl, pH 4.1-buffer solution containing iron at a final concentration of 0.6 or 0.3 mg/ml after PD-10 purification (Fe.sub.3O.sub.4 particles coated with catechol-PEG400 and carboxy group).

[0099] The liposome solution was freeze-thawed 3 times and unilamellar liposomes were prepared with a needle tip sonicator (515 sec low energy on ice). The buffer was changed to PBS pH 6.5 by using a PD-10 column. Both doxorubicin (1.4 mM doxorubicin for 10 M lipids) and liposomes were pre-heated to 50 C. if DPPC was used and to 60 C. if DSPC was used and then combined and incubated for 30 min at 50 C. or 60 C. in a water bath. Liposomes were purified from the unbound compounds using a PD-10 column. Elution curves were established by using doxorubicin absorbance at 485 nm. Liposomes were controlled by liposome size measurements using dynamic light scattering (DLS).

[0100] b) Cisplatin Loaded Liposomes:

Liposomes were made using a lipid mixture of DSPC/cholesterol/SM/DOTAP (20:30:30:20 mol %, respectively) consisting of a total of 20 M of lipids. Lipids stored in chloroform were pipetted to a round bottomed flask, dried under nitrogen and lyophilized for 2 h to remove trace amounts of chloroform. Lipids were allowed to hydrate for 30 min in 60 C. buffered water solutions containing cisplatin (9 mM) and nanoparticle iron at a final concentration of 0.6 or 0.3 mg/ml after PD-10 purification (Fe.sub.3O.sub.4 particles coated with catechol-PEG400 and carboxy group). The liposome solution was freeze-thawed 10 times. Extrusion was performed 11 times through a 100 nm polycarbon membrane by using a small volume extruder, or unilamellar liposomes were prepared using a needle tip sonicator (415 sec low energy on ice). Liposomes were purified from the unbound compounds using a PD-10 column. Elution curves were obtained using cisplatin absorbance (at 301 nm in 0.1 M HCl). Liposomes were controlled by liposome size measurements using DLS.

[0101] c) Paclitaxel Loaded Liposomes:

Liposomes were made using a lipid mixture of DSPC/cholesterol/SM/DOTAP (20:30:30:20 mol %, respectively) consisting of a total of 20 M of lipids. Lipids stored in chloroform were pipetted to a round bottomed flask, dried under nitrogen and lyophilized for 2 h to remove trace amounts of chloroform. Paclitaxel was added to the lipids (0.7 to 1 mg) and traces of chloroform were lyophilized overnight. Lipids and paclitaxel were allowed to hydrate for 30 min at 60 C. in buffered water solutions containing Fe.sub.3O.sub.4 nanoparticles (0.6 mg/ml or 0.3 mg/ml final concentration). Liposome solutions were freeze-thawed 7-10 times. Extrusion was performed 11 times through a 100 nm polycarbon membrane using a small volume extruder, or unilamellar liposomes were prepared using a needle tip sonicator (515 sec low energy on ice). Liposomes were purified from the unbound compounds using a PD-10 column. Elution curves were obtained using paclitaxel absorbance (227 nm). Liposomes were controlled by liposome size measurements using DLS.

[0102] d) Indocyanine Green (ICG) Loaded Liposomes:

Liposomes were made using a lipid mixture consisting of a total of 20 M of lipids (DSPC/Cholesterol/SM/DOTAP (20:30:30:20) in mol/mol ratios). Lipids stored in chloroform were pipetted to a round bottomed flask, dried under nitrogen and lyophilized for 2 h to remove trace amounts of chloroform. Lipids were allowed to hydrate for 30 min at 60 C. in buffered water solutions containing ICG (0.5 mg/ml) and Fe.sub.3O.sub.4 nanoparticles in final concentration 0.6 or 0.3 mg/ml after PD-10 purification (Fe.sub.3O.sub.4 particles coated with catechol-PEG400 and carboxy group). Extrusion was performed 11 times through a 100 nm polycarbon membrane using a small volume extruder, or unilamellar liposomes were prepared using a needle tip sonicator (315 sec low energy on ice). Liposomes were purified from the unbound compounds using a PD-10 column. Elution curves were obtained using ICG absorbance at 800 nm. Liposomes were controlled by liposome size measurements using DLS.

Example 2

AMF Treatment

[0103] The inventors used a coil with a magnetic field strength of 1.5 mT for in vivo experiments and the same 1.5 mT coil as well as a device with variable field strength ranging from 5-23 mT for in vitro studies.

[0104] AMF treatments were performed for 20 min (100 kHz, 15V, 1.5 mT) for each sample containing 300-500 l liposomes. If a sample was pretreated, SMase 0.4 U were added (final activity 0.77 U/ml) 30 minutes before applying the AMF treatment. After treatment, samples were inserted into a snake skin dialysis bag and dialyzed for 24 h against PBS, or PBS pH 6.5 in order to analyze the amount of free doxorubicin, cisplatin, paclitaxel or free ICG. After 24 h incubation both solutions inside and outside of the bag were analyzed by measuring absorbance of doxorubicin, cisplatin, paclitaxel or ICG and the leakage rate was calculated.

[0105] AMF was also generated using instrumentation from nanoScale Biomagnetics (nB, Zaragoza, Spain). In this case, fields between 5-23 mT and 229-823 kHz were used. Different concentrations (0.15 to 1.2 mg/ml of Fe.sub.3O.sub.4 nanoparticles per 13 mol/ml lipids in PBS) and also different sizes (5 to 20 nm) of Fe.sub.3O.sub.4 nanoparticles were incubated for 10 min in 5-23 mT in 750 l, to study both leakage of the liposomes and magnetic hyperthermia, as well as to analyze specific power absorption values.

Example 3

aSMase Detection

[0106] aSMase detection was performed with the aSMase detection kit by Cayman Chemicals (USA) according to the instructions of the manufacturer. aSMase levels were measured 10 minutes after irradiation with 0 to 20 Gy gamma radiation.

Example 4

In Vivo Imaging

[0107] In vivo imaging experiments were done with athymic nude mice bearing orthotopic SCC-4 carcinoma in the lower jaw region. Animals were anesthetized with Ketamine (110 mg/kg bw)/Xylazine (16 mg/kg bw) during handling and imaging.

[0108] After purification, liposomes containing ICG fluorophores and Fe.sub.3O.sub.4 nanoparticles in a volume of 150 l were administered by tail vein injection. The mice were imaged with a Berthold NightOWL camera (Berthold Technologies, Bad Wildbad). Image analysis was performed with indigo software. Mice were also imaged with a fluorescent tomograph (FMT 2500, Perkin Elmer, Inc., Waltham, Mass., USA) and under a surgical microscope (Mller-Wedel, Wedel, Germany).

[0109] Mice were positioned so that the tumor was inside the magnetic coil of the AMF apparatus. AMF (1.5 mT and 100 kHz) was applied to the tumor for 20 min. Mice were imaged after one hour together with the control mice that had gotten the same liposomal injection but had not been submitted to AMF. A clear increase in the ICG signal due to leaked ICG showed on the tumor area under the chin of the mouse.

[0110] Experts in this field will appreciate that the invention described in these examples is open to variations and modifications other than those specifically described. It is to be understood that the invention includes all such variations and modifications.

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