Laser activated nanothermolysis of cells

Provided herein are methods and systems to increase selective thermomechanical damage to a biological body, such as a cancer cell or cell associated with a pathophysiological condition. The biological body or cancer cell is specifically targeted with nanoparticulates comprising one or more targeting moieties which form nanoparticulate clusters thereon or therewithin. Pulsed electromagnetic radiation, e.g., optical radiation, having a wavelength spectrum selected for a peak wavelength near to or matching a peak absorption wavelength of the nanoparticulates selectively heats the nanoparticulates thereby generating vapor microbubbles around the clusters causing damage to the targets without affecting any surrounding medium or normal cells or tissues. Also provided are methods for treating leukemia and for selectively and thermomechanically causing damage to cells associated with a pathophysiological condition using the methods and system described herein.

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Description
CROSS-REFERENCE TO RELATED APPLICATIONS

This is a continuation-in-part application under 35 U.S.C. §120 of pending U.S. Ser. No. 11/795,856, filed Jul. 23, 2007, which is a national stage application under 35 U.S.C. 371 of international application PCT/US2006/002186, filed Jan. 22, 2006, now abandoned, which claims benefit of priority under 35 U.S.C. 119(e) of provisional U.S. Ser. No. 60/646,018, filed Jan. 22, 2005, now abandoned, the entirety of all of which are hereby incorporated by reference.

BACKGROUND OF THE INVENTION

1. Field of the Invention

The present invention relates to the fields of medical therapies employing electromagnetic radiation and of nanoparticles. More specifically, the present invention relates to a method and a system for electromagnetic radiation induced selective destruction of abnormal biological bodies or structures utilizing bioconjugated nanoparticles.

2. Description of the Related Art

In a variety of medical applications, it is desirable to inactivate, ablate or otherwise achieve the destruction and elimination of abnormal cells, while preserving normal and healthy cells. Recent advances in biomedical optics and nanotechnology has created a solid basis for development of effective therapies for human diseases, such as cancer or atherosclerosis. The major breakthrough in this area is the possibility for a therapeutic agent to target selectively certain types of cells with molecular specificity. Furthermore, new imaging modalities are being developed to visualize not only abnormal cells and tissues, but also to monitor and guide therapeutic procedures, making them more effective and safe. The prior art discloses laser (optical) methods of therapy, which can be, in principle, guided by imaging based on optical contrast or otherwise enabled by optical activation.

Selective and specific inactivation, that is, damage or destruction, of cells with optical means requires substantial optical contrast between target cells and all non-target normal cells. General approaches to achieve cell inactivation include thermomechanical damage through high-intensity pulsed laser interaction with cells, thermal damage through continuous wave high power interaction, such as used for hyperthermia and coagulation, and biochemical damage through relatively low power interactions of photons with molecules resulting in production of chemically active species, such as ions, radicals and metastable excited states (1). Similarly, these types of interactions can be the result of interactions of electromagnetic radiation of various wavelengths, e.g., X-rays, UV, visible light, near-infrared, infrared photons, microwave, and radiofrequency quanta, and also high intensity acoustic waves.

Laser ablation is based on thermal and mechanical effects caused in cells by absorption of high-intensity laser pulses (2). One of the advanced concepts of laser ablation, selective photothermolysis, was introduced more than 20 years ago (3). Selective photothermolysis employs short laser pulsed interaction with absorbing tissue microstructures to induce localized physical damage by avoiding thermal diffusion of the deposited laser energy. Other advanced concepts of precise laser ablation with limited thermal collateral damage to surrounding tissue employs conditions of pressure confinement upon short pulse laser irradiation to produce tensile wave causing cavitation bubbles at temperatures below 100° C. (4) or ultrashort laser pulses to cause nonlinear absorption and rapid micro-explosion that occurs prior to thermal diffusion (5). Slow heating with long pulses of electromagnetic radiation can also be used for selective damage of cells, however, due to thermal diffusion requires much stronger optical contrast to achieve specificity and to not damage adjacent cells or tissues (6). As an alternative to instantaneous thermomechanical destruction of target cells, photodynamic therapy employs low-intensity laser irradiation to produce necrosis and apoptosis through delayed damage mechanisms caused by photochemically produced toxic species, for example, radicals and singlet oxygen (7).

Both high and low intensity interactions of laser radiation with cells can be made selective and specific to target cells. Such selectivity requires utilization of either endogenous or exogenous chromophores, which strongly absorb specific colors of laser spectrum in the visible and near-infrared, i.e. in the range of wavelengths where majority of tissue constituents do not absorb. Since endogenous tissue chromophores strongly absorbing in the red and near-infrared are limited to hemoglobin and melanin, the medical applications of natural contrast agents are limited to blood vessels and retinal pigmented epithelium (8,9). Selective damage of leukemia and other cancerous cells demands exogenous contrast agents. Optical contrast agents that possess very strong absorption of near-infrared radiation attract attention of researchers because normal cells and tissues are transparent in this spectral range, so there is a great potential to achieve an exceptional selectivity of cell damage through the contrast agent. Chromophore-assisted laser inactivation (CALI) is a term to describe selective inactivation of certain proteins in cellular membranes using laser irradiation of cells stained with molecular dyes strongly absorbing red laser pulses (10).

It was recently revealed that gold nanoparticles can be designed to absorb any desirable color of near-infrared radiation by either changing the thickness gold shells on the silica core (11) or changing the aspect ratio of gold nanorods, i.e., ellipsoids or other prisms with one elongated axis (12). Gold nanoparticles and especially silver nanoparticles absorb near-infrared light much stronger than nanoparticles of organic dyes, which makes them superior contrast agents for imaging a small cluster of cancer cells in the depth of tissue (13). Nanoparticle assisted selective laser thermolysis of cells was recently demonstrated by targeting optically absorbing nanoparticles to cell surface receptors and superheating them with laser pulses (14). Based on the experimental results obtained with microparticles, the prior art speculated that the cavitation bubble generation that results in cell inactivation after laser irradiation with certain optical fluence may depend on particle size. On the other hand, the prior art neither provides an explanation of the underlying physical phenomena, nor provides a solution for achieving highly effective cell damage using low threshold fluence of electromagnetic radiation.

U.S. Pat. No. 6,530,944 teaches optically active nanoparticles that can be used in therapeutic and diagnostic methods. However, therapeutic applications disclosed by West et al. are limited to methods of hyperthermia, i.e. slow heating usually with continuous wave lasers and other optical sources. Nanoparticles will extravasate preferentially at locations where the blood vessel walls have increased porosity or have microvascular surface changes, especially at tumor sites. O'Neal et al. demonstrated that intravenous injection of gold nanoshells, which are nanoparticles strongly absorbing in the near-infrared spectral range, into a mouse, resulted in nonspecific, but effective targeting of an implanted tumor. Further, it enabled hyperthermic damage of the tumor through heating the nanoshells to temperatures several degrees higher than that in surrounding tissue via continuous wave laser illumination of the tumor area (15). U.S. Pat. No. 6,165,440 taught that a combination of radiation and nanoparticles or microparticles could be used to temporarily open pores in cancer cell membranes and blood vessels to allow better penetration of drugs into solid tumors.

U.S. Pat. Nos. 6,530,944 and 6,699,724 disclosed optical diagnostic uses of nanoparticles that emit near-infrared light, i.e., nanodots or absorb near-infrared red light, i.e., nanoshells. Sokolov et al. utilized the capability of gold nanoparticles to reflect light strongly and thereby enhance contrast of abnormal or cancerous tissue specifically targeted with nanoparticles conjugated with monoclonal antibody (16). Oraevsky et al. (17,18) predicted that various nanoparticles can enhance optical absorption in tissue and emit thermoacoustic waves, which in turn can be utilized in optoacoustic imaging of tissue. U.S. Patent Pub. No. 20050175540 described non-spherical nanoparticles and a method by which to optoacoustically detect the presence of objects as small as 1-mm in a body, which can be penetrated by electromagnetic radiation. It was discovered that at least partially metallic nanoparticulates fabricated or manipulated to be non-spherical not only will shift the optical absorption spectrum into the near-infrared range for deeper penetration of radiation into a body, but also will both narrow the absorption band and simultaneously increase the effective absorbance, in certain instances by more than an order of magnitude. This greatly increases the optoacoustic efficacy of the nanoparticulate, making the manipulated nanoparticulate a very high contrast optoacoustic imaging agent.

U.S. Pat. Nos. 5,213,788, 5,411,730, 5,427,767, 5,521,289, 6,048515, 6,068,857, 6,165,440, 6,180,415, 6,344,272, 6,423,056, and 6,428,811 disclose various electromagnetically active nanoparticles for use as therapeutic or imaging contrast agents. It was established by many groups that specific targeting using antibodies increases efficacy of anticancer therapies (19). In addition to IgG type antibodies, short peptides can be used as targeting vectors (20).

It is well recognized that any minimally invasive therapy may benefit from imaging methods that can precisely guide the therapeutic procedure. Lapotko et al. have developed methods of photothermal detection and imaging that enable one to visualize individual cells and thermomechanical processes that occur in cells upon pulsed laser irradiation (21-22). U.S. Pat. Nos. 5,840,023 and 6,309,352 taught a method and a system of optoacoustic imaging that helps detection, localization and real-time monitoring of abnormal tissue in the depth of normal tissue.

Chemotherapy and radiotherapy are often ineffective in treating human cancer, including hematological malignancies, such as leukemia. The state of the art therapy methods and systems have serious limitations, associated with significant treatment toxicity and generation of drug-resistant tumor cells (24-27). The residual cells, therefore, must be eliminated from blood or bone marrow grafts by methods generally called “purging”. Available purging methods employ pharmacological and photochemical (PDT) treatment, magnetic and fluorescence based sorting and elimination in cuvette through adsorption to mab attached to its bottom. These methods provide help and relief to cancer patients, but do not provide sufficient efficacy, i.e. 100% elimination, or adequate speed of cell elimination (28). New treatment strategies that are more effective, faster and less expensive are therefore necessary to overcome these problems.

However, there is a recognized need in the art for an effective and safe method and system that provides selective lysis, i.e., destruction, of targeted abnormal cells and other microstructures or microbodies while leaving all normal cells intact, using electromagnetic radiation. Specifically, the prior art is deficient in therapeutic laser methods and systems utilizing nanoparticulate contrast agents. More specifically, the prior art is deficient in methods and systems of Laser Activated Nano-Thermolysis Cell Elimination Technology (LANTCET) that utilize metal nanoparticles for laser therapy of cancer. The present invention fulfils this longstanding need in the art.

SUMMARY OF THE INVENTION

The present invention is directed to a method for increasing selective therapeutic thermomechanically induced damage to a biological body. The method comprises specifically targeting a biological body comprising a medium with a plurality of nanoparticulates each conjugated to at least one targeting moiety. The nanoparticulates are effective to form one or more nanoparticulate clusters on or in the biological body upon targeting thereto. The biological body is irradiated with at least one pulse of electromagnetic radiation having a spectrum of wavelengths selected to have a peak wavelength that is near to or that matches a peak absorption wavelength of the nanoparticulates. Subsequently, vapor microbubbles are generated from heat produced via absorption of the electromagnetic radiation into the nanoparticulates such that the microbubbles cause selective and increased thermomechanical damage to the targeted biological body. In a related invention the method further may comprise filtering the products of the thermomechanical damage from the medium. In another related invention the method further may comprise receiving a photothermal signal or generating an optical image of thermomechanical effects to monitor and to guide selective thermomechanical damage to the biological body.

The present invention also is directed to a system for increasing selective therapeutic thermomechanical damage to abnormal cells. The system comprises a chamber containing the abnormal cells in a medium, a source of nanoparticulates adapted to specifically target the abnormal cells which is fluidly connected to the cell chamber, an optical chamber adapted to contain the targeted abnormal cells which is fluidly connected to the cell chamber, and a means for filtering out cells damaged by thermomechanical effects resulting from absorption of the electromagnetic radiation emitted at the peak wavelength which is fluidly connected to the cell chamber. A pulsed source of electromagnetic radiation is directed against the targeted cancer cells in the optical chamber, where the source is configured to emit a spectrum of wavelengths selected to have a peak wavelength that is near to or that matches a peak absorption wavelength of the nanoparticulates. In a related invention the system further comprises a means for receiving a photothermal signal or for generating an optical image of the thermomechanical effects.

The present invention is directed further to a method for treating a leukemia in an individual. The method comprises obtaining a sample comprising normal and leukemic cells from the individual and placing the sample in the cell chamber of the system described herein. The cancer cells in the sample are targeted with the nanoparticulates described herein and the targeted cancer cells are irradiated with electromagnetic radiation emitted from the pulsed source comprising the system. The electromagnetic radiation is absorbed by the nanoparticulates thereby causing selective and increased thermomechanical effects damaging to the targeted cancer cells, but not to the normal cells comprising the sample. The damaged cells are filtered out from the sample and the normal cells remaining in the sample are returned to the individual thereby treating the leukemia. The method steps may be repeated zero or more times. In a related invention the method further may comprise receiving a photothermal signal or generating an optical image of thermomechanical effects to monitor and guide selective thermomechanical damage to the biological body.

The present invention is directed further still to a method for selectively and thermomechanically damaging cells associated with a pathophysiological condition. The method comprises targeting the cells with a first monoclonal antibody specific thereto and targeting the cells with gold nanoparticulates described herein that are modified with a second monoclonal antibody specific to the first monoclonal antibody whereupon one or more clusters of the nanoparticulates form on or in the targeted cells. One or more of the clusters of gold nanoparticulates formed on or in the targeted cells are heated and vapor microbubbles are generated around the heated clusters sufficient to thermomechanically damage the cells. In a related invention, the method further comprises photothermally or optically monitoring the thermomechanical damage.

Other and further aspects, features, benefits, and advantages of the present invention will be apparent from the following description of the presently preferred embodiments of the invention given for the purpose of disclosure.

BRIEF DESCRIPTION OF THE DRAWINGS

The appended drawings have been included herein so that the above-recited features, advantages and objects of the invention will become clear and can be understood in detail. These drawings form a part of the specification. It is to be noted, however, that the appended drawings illustrate preferred embodiments of the invention and should not be considered to limit the scope of the invention.

FIGS. 1A-1B show the calculated optical absorption spectra of gold (FIG. 1A) and of silver (FIG. 1B) nanorods. With an increasing aspect ratio, the peak of plasmon resonance absorption gradually shifts to longer wavelengths in the near-infrared. Aspect ratios left to right in graphs: 2a=20 nm, 2c=60 nm; 2a=20 nm, 2c=100 nm; 2a=20 nm, 2c=140 nm.

FIGS. 2A-2C show electron microphotographs (FIGS. 2A-2B) of gold nanorods with diameter of 15 nm and 2 different aspect ratios, 3.2 and 6.9, placed on the surface of a glass slide, and an experimentally measured optical absorption spectra (FIG. 2C) of these gold nanorods, conjugated with PEG and suspended in water.

FIG. 3 shows optoacoustic signal as a function of laser fluence incident upon suspension of spherical gold nanoparticles with 100 nm diameter. Absorption cross-section of these nanoparticles is σα=1.4×10−9 cm2.

FIG. 4 shows threshold fluence for the laser damage of tumor cells targeted with gold nanoparticles that formed clusters inside the target cells.

FIG. 5 is a sketch of a LANTCET system modified from a blood dialysis system.

FIGS. 6A-6B comprise a schematic diagram of LANTCET depicting simultaneous monitoring and guidance by photothermal microscopy.

FIGS. 7A-7B are electron microscopy images of K562 cells. FIG. 7A is a control cell without nanoparticles and FIG. 7B is a cell specifically targeted with primary MABs (CD 15 and glycophorin A) and 30 nm diameter spherical gold nanoparticles conjugated to secondary MAB. Small black dots are single nanoparticles, larger black spots are clusters of nanoparticles.

FIGS. 8A-8C illustrate PT responses obtained after a single laser pulse at 532-nm wavelength and duration of 10-ns. FIG. 3A shows a single K562 cell targeted with bare nanoparticles—no bubble, no damage at optical fluence of 35 J/cm2. FIG. 8B shows a suspension of 30-nm diameter single gold nanoparticles; optical fluence of 35 J/cm2. FIG. 8C shows a single K562 specifically targeted cell with clusters of NP, optical fluence of 5 J/cm2.

FIG. 9 illustrates the cell damage probabilities, DP, experimentally obtained for different pump laser pulse (532 nm, 10 ns) fluence levels for specifically targeted common B acute lymphoblasts (lymphoblasts) and normal stem cells, for specifically targeted K562 cells (test), for non-specifically targeted K562 cells (control #1), for K562 cells targeted with bare NP (control #2) and for untargeted K562 cells (control #3).

FIGS. 10A-10B are optical microscopic images of normal stem cells (FIG. 10A) and common B acute lymphoblasts (FIG. 10B) in the cuvette after irradiation of cell suspension with a single broad laser pulse with optical fluence of 1.7 J/cm2. Cells were prepared using specific targeting protocol.

FIGS. 11A-11B are schematic diagrams of cell irradiation and laser photothermal microscopy for a single cell (FIG. 11A) and for a cell suspension (FIG. 11B) of ALL tumor cells and normal cells.

FIGS. 12A-12F are optical (FIGS. 12A-12B) and fluorescent (FIGS. 12C-12D) images of human leukemia cells (B-lymphoblasts) and corresponding fluorescence signal profiles (FIGS. 12E-12F) for images obtained after the stage 1, incubation at 48 C (FIGS. 12A, 12C, 12E) and after the stage 2, incubation at 378 C (FIGS. 12B, 12D, 12F). White lines on the cell images show the cross-section for amplitude profile and cell boundaries (nuclei and of outer membranes).

FIGS. 13A-13B are histograms of fluorescent image parameters of tumor cells: the maximal values of pixel amplitude in peaks Max (FIG. 13A) and radial distribution Mir of fluorescent peaks (FIG. 13B).

FIGS. 14A-14B illustrate the kinetics of the clusterization of NPs during cell incubation at 48° C. (stage 1) and 37° C. (stage 2) for normal BM cells (FIG. 14A) and tumor cells (FIG. 14B) (O-37° C., peaks at membrane, -37° C., peaks inside cell, Δ-4° C., peaks at membrane, ▴4° C., peaks inside cell).

FIGS. 15A-15C illustrate a typical bubble-specific PT response (FIG. 15A), PT-image (FIG. 15B), and optical image of same cell before irradiation (FIG. 15C) obtained for a tumor cell after single laser pulse for the incubation conditions being 37° C., 2 hours.

FIGS. 16A-16B illustrate the influence of incubation parameters on cell damage. FIG. 16A shows the level of survived live tumor cells LLC experimentally obtained after a single laser pulse (532 nm, 0.6 J/cm2) for primary MABs CD19, CD20 and CD 22 applied during the first stage of cell targeting with spherical nanoparticles. FIG. 16B shows the level of survived live cells LLC experimentally obtained for tumor CD10+ cells and normal BM cells after their irradiation with single laser pulse (532 nm, 0.6 J/cm2) at different temperatures during the second stage of incubation.

FIGS. 17A-17C are images and spectra of K562 myeloid culture cells obtained with an optical scattering microscope. FIG. 17A shows a cell without nanorods. FIG. 17B shows a cell with cytoplasm-located clusters of gold nanorods. FIG. 17C is an optical scattering spectra for intact cells (red) and for nanorod clusters in the cell (black).

FIGS. 18A-18D are optical microscopic images of an AML cell before (FIG. 18A) and 5 sec after (FIG. 18D) cell treatment with single laser pulse (10 ns, 780 nm and a photothermal image (FIG. 18B) and response (FIG. 18C) of laser-induce PTB in the cell shown in FIG. 18B.

FIGS. 19A-19B illustrate bubble generation probability spectra (FIG. 19A) and bubble lifetime spectra (FIG. 19B) for single gold nanorods and human bone marrow tumor cells.

FIGS. 20A-20D are optical microscopic images of Hep-2C cell before (FIG. 20A) and 5 sec after (FIG. 20D) cell treatment with single laser pulse (10 ns, 720 nm) and a photothermal image (FIG. 20B) and response (FIG. 20C) of laser-induce PTB in the cell shown in FIG. 20B.

FIGS. 21A-21B are vertical cross sections of the tumor that was treated with spherical nanoparticles and laser pulse (FIG. 21A) and the control tumor which was not treated with spherical nanoparticles but was treated with the laser (FIG. 21B). Blue is an intact area and white is a necrotic area, scale bar in mm.

 DETAILED DESCRIPTION OF THE INVENTION

In one embodiment of the present invention there is provided a method for increasing selective therapeutic thermomechanically induced damage to a biological body, comprising specifically targeting a biological body comprising a medium with a plurality of nanoparticulates each conjugated to at least one targeting moiety, where the nanoparticulates are effective to form one or more nanoparticulate clusters on or in the biological body upon targeting thereto; irradiating the biological body with at least one pulse of electromagnetic radiation having a spectrum of wavelengths selected to have a peak wavelength that is near to or that matches a peak absorption wavelength of the nanoparticulates; and generating vapor microbubbles from heat produced via absorption of the electromagnetic radiation into the nanoparticulates where the vapor microbubbles cause selective and increased thermomechanical damage to the targeted biological body.

Further to this embodiment the method may comprise filtering the products of the thermomechanical damage from the medium. In another further embodiment the method may comprise receiving a photothermal signal or generating an optical image of thermomechanical effects to monitor and guide selective thermomechanical damage to the biological body. In all embodiments the biological body may be an abnormal cell, a bacterium or a virus.

In these embodiments the nanoparticulates may have a dimension of about 1 nm to about 1000 nm. Also, the nanoparticulate cluster may have a total volume about 2 to about 200 times greater than a volume of the nanoparticulate comprising the same. In addition, the nanoparticulate may be a spherical nanoparticle, a nanorod or a nanoshell at least partially comprising gold or silver or is a carbon nanotube.

Furthermore, in all embodiments the nanoparticulates comprise a targeting moiety that is a monoclonal antibody or a peptide specifically targeted to a receptor site on the biological body. In an aspect, the receptor site further comprises another monoclonal antibody or peptide attached thereto specific for the targeted monoclonal antibody. Further to all embodiments the nanoparticulates may comprise complementary strands of a nucleic acid conjugated thereto or a combination thereof. Further still the nanoparticulate may comprise PEG molecules.

In all embodiments the wavelength spectrum of the pulse of electromagnetic radiation may have a range of wavelengths of about 300 nm to about 300 mm. In an aspect the pulse of electricalmagnetic radiation is optical radiation. This optical radiation may have a wavelength in the range from 500 nm to 1150 nm. Also, in all embodiments the pulse of electromagnetic radiation is about 1 ns to about 100 ns.

In another embodiment of the present invention there is provided a system for increasing selective therapeutic thermomechanical damage to abnormal cells, comprising a chamber containing the abnormal cells in a medium; a source of nanoparticulates modified to specifically target the abnormal cells fluidly connected to the cell chamber; an optical chamber adapted to contain the targeted cancer cells fluidly connected to the cell chamber; a pulsed source of electromagnetic radiation directed against the targeted cancer cells in the optical chamber, where the source is configured to emit a spectrum of wavelengths selected to have a peak wavelength that is near to or that matches a peak absorption wavelength of the nanoparticulates; and means for filtering out cells damaged by thermomechanical effects resulting from absorption of the electromagnetic radiation emitted at the peak wavelength, where the filtering means is fluidly connected to the cell chamber. In a further embodiment the system comprises a means for receiving a photothermal signal or for generating an optical image of the thermomechanical effects.

In these embodiments the nanoparticulates each comprise at least one targeting moiety specifically targeted to a receptor site on the cancer cell. Furthermore, the receptor site may comprise another targeting moiety attached thereto which is specific for the targeting moiety on the nanoparticulates. Examples of a targeting moiety are a monoclonal antibody or a peptide. In another further embodiment the nanoparticulates further may comprise complementary strands of a nucleic acid conjugated thereto or a combination thereof. Further still the nanoparticulate may comprise PEG molecules.

In all embodiments the abnormal cells may be leukemic cancer cells. Also, in all embodiments the dimensions, shapes, or metal or carbon compositions of the nanoparticulates or nanoparticulate clusters are as described supra. Furthermore, the spectrum wavelengths and types of electromagnetic radiation and time of pulse duration are as described supra.

In yet another embodiment of the present invention there is provided a method for treating a leukemia in an individual, comprising a) obtaining a sample comprising normal and leukemic cells from the individual; b) placing the sample in the cell chamber of the system described supra; c) targeting the leukemic cells in the sample with the modified nanoparticulates comprising the system; d) irradiating the targeted leukemic cells with electromagnetic radiation emitted from the pulsed source comprising the system, where the electromagnetic radiation absorbed by the nanoparticulates causes selective and increased thermomechanical effects damaging to the targeted cancer cells, but not to the normal cells comprising the sample; e) filtering out the damaged cells from the sample; f) returning the normal cells remaining in the sample to the individual; and g) repeating the method steps a) to f) zero or more times, thereby treating the leukemia.

Further to this embodiment the method comprises receiving a photothermal signal or generating an optical image of the thermomechanical effects to monitor and guide selective thermomechanical damage to the cancer cells. In both embodiments the thermomechanical effects are caused by heat generated within the nanoparticulates from absorbed electromagnetic radiation sufficient to form vapor microbubbles around the clusters.

In yet another embodiment of the present invention there is provided a method for selectively and thermomechanically damaging cells associated with a pathophysiological condition, comprising targeting the cells with a first monoclonal antibody specific thereto; targeting the cells with gold nanoparticulates modified with a second monoclonal antibody specific to the first monoclonal antibody whereupon one or more clusters of the nanoparticulates form on or in the targeted cells; heating one or more clusters of gold nanoparticulates formed on or in the targeted cells; and generating vapor bubbles around the heated clusters sufficient to thermomechanically damage the cells. Further to this embodiment, the method may comprise photothermally or optically monitoring the thermomechanical damage. In both embodiments, the gold nanoparticulates and nanoparticulate clusters and the optical radiation are as described supra. Also, in both embodiments, the cells associated with a pathophysiological condition may be a cancer cell, such as a leukemic or solid tumor cancer cell, a bacterial cell or a virus.

As used herein, the term “a” or “an”, when used in conjunction with the term “comprising” in the claims and/or the specification, may refer to “one,” but it is also consistent with the meaning of “one or more,” “at least one,” and “one or more than one.” Some embodiments of the invention may consist of or consist essentially of one or more elements, method steps, and/or methods of the invention. It is contemplated that any method or composition described herein can be implemented with respect to any other method or composition described herein.

As used herein, the term “or” in the claims refers to “and/or” unless explicitly indicated to refer to alternatives only or the alternatives are mutually exclusive, although the disclosure supports a definition that refers to only alternatives and “and/or.”

As used herein, the term “nanothermolysis” refers to damage, ablation, destruction of biological target cells assisted and enabled by nanoparticles.

As used herein, the phrase “Laser Activated Nano-Thermolysis” refers to selective damage of cells using nanoparticles (nanoparticles) targeted to specific receptors expressed in cancer cells, but not in normal cells.

As used herein, the phrase “cell elimination” refers to a successful treatment procedure. Thus, as used herein, the phrase “Laser Activated NanoThermolysis Cell Elimination Technology” or “LANTCET” refers to a method and a system for electromagnetic radiation induced selective destruction of abnormal biological structures utilizing bioconjugated nanoparticles, such as tissues, cells, bacteria and viruses. Because of their selectivity for abnormal biological structures, LANTCET may be applied repeatedly to insure or achieve a desired level of cell elimination, up to and including complete cell elimination, although a single LANTCET application with zero repeats is well within the scope of the invention.

As used herein, the term “nanoparticulate” refers to a single nanoparticle, a collection of nanoparticles, or a nanoparticle aggregate. As used herein, the term “nanoparticulate cluster” describes specific aggregation of nanoparticles in which nanoparticles may touch each other or be in proximity to each other, so that when irradiated with laser pulses or other pulses of electromagnetic energy, they present themselves as a single unresolved source of thermal energy.

As used herein, the phrase “at least partially metallic” refers to a preferred nanoparticulate effective to absorb electromagnetic radiation by the plasmon resonance mechanism, which is known to yield very strong absorption coefficient.

Provided herein are methods and systems using Laser Activated Nano-Thermolysis Cell Elimination Technology or LANTCET for the electromagnetic radiation induced selective destruction and purging of abnormal biological structures, such as tissues, cells, bacteria and viruses. LANTCET has advantages over current methods of purging as shown in Table 1.

TABLE 1 Magnetic Flow Method separators Cytometry PDT LANTCET Rate, cell/s 107 104 107  107  Volume, ml 50-100 50   >1000 >1000 per one cycle Max amount of 109 106 1010 1010 treated cells Efficacy, % of 90 99.9   95    99.9 labeled cells Safety Additional Potential Poten- Potential chemical cell tially laser treatment of the damage toxic for damage to cells for removing at high normal 0.1-3% of magnetic particles rate cells normal cells may damage normal cells

Although magnetic elimination is not damaging to normal cells, its main disadvantage is that efficacy of this method is not sufficient for purging residual cancer cells due to a lack of strong magnetic force that separates small ferrous beads coupled to the target cells from normal untargeted cells. The main advantage of flow cytometry is its high efficiency of cell sorting, however, the low throughput of this method is a limiting factor. Photodynamic therapy lacks sufficient specificity to target cells due to low contrast provided by organic dyes. Overall, these methods and systems lack means for high through-put that is effective, specific for the target cells and safe.

In a LANTCET method, strongly absorbing nanoparticles provide contrast relative to untargeted cells that is unmatched by any other contrast agent. Furthermore, formation of clusters of nanoparticles in cells makes laser thermolysis of the target cells by microbubbles a very effective, low threshold process that is harmless to normal untargeted cells. Selectivity is conferred by targeting an abnormal biological structure with bioconjugated nanoparticles. These nanoparticles are targeted to molecular receptors on the surface of the target cells using antibodies, such as immunoglobulin type proteins, or short peptides, which help to achieve selectivity of nanoparticle accumulation in cells.

The targeting protocol is designed to produce clusters of nanoparticles on the surface of cells and/or inside the abnormal cells in order to enhance efficacy of thermomechanical interactions of electromagnetic radiation with the nanoparticles. Such enhancement results in a substantially lower threshold of the fluence or the power required to achieve the abnormal cell damage. This in turn increases specificity and probability of damage to the targeted cell damage and of safety to normal cells and tissues. The methods and systems described herein have a variety of applications including, but not limited to, minimally invasive therapy of cancer, such as eliminating early subsurface tumors from tissue and purging leukemia cells from bone marrow graft transplants or blood.

Generally, the methods and systems described herein are effective to substantially increase the thermomechanical damage to an abnormal cell, biological body or microbody or microstructure, such as, but not limited to, a biological cell, e.g., a cancer or tumor cell, a bacterium or other microorganism, a virus, or atherosclerotic plaques, resulting from a pulse of electromagnetic radiation absorbed by the nanoparticles and to decrease the damage to the surrounding normal body. The abnormal cell, biologic body, cell or virus may be associated with a pathophysiological condition, such as, but not limited to a cancer or other malignancy, a bacterial or viral infection or atherosclerosis. More particularly, the methods and systems described herein include the following components and steps.

Nanoparticles

The nanoparticulates provided herein are designed with the maximum capability to absorb electromagnetic radiation of a wavelength effective to penetrate the body, but not to be absorbed by molecular content of the body. That is, the nanoparticulates must absorb electromagnetic radiation very strongly, so that such absorption is sufficient to superheat them well above the boiling point of the surrounding biological medium, e.g., water or a water like medium, with a fluence of energy that produces insignificant direct heating of the surrounding medium. Nanoparticulates that strongly absorb electromagnetic radiation in the near-infrared spectral range, the range of wavelengths where molecular constituents of biological cells and tissues possess no or minimum absorption are used. Preferably, the nanoparticles may be at least partially metal nanoshells, metal nanorods or carbon nanotubes.

One skilled in the art can predict that many combinations and complex structures can be created based on basic properties of nanoparticles known in the art. It is contemplated that clusters of other nanoparticulates, such as carbon nanotubes and liposomes filled with organic dyes, can be utilized for LANTCET since clusters of these particles can strongly absorb near-infrared radiation. Although, the near-infrared spectral range seems to be the preferred range based on what is known in the art, nanoparticles may be designed to absorb x-rays, microwave (RF) radiation or visible radiation. The absolute value of the absorption coefficient for a given wavelength of electromagnetic radiation is not as important as the contrast, that is, difference, ratio between the absorption coefficient in the nanoparticulate and the background surrounding medium or body that was not targeted with said nanoparticulate.

The most preferred materials for nanoparticulate composition are gold and silver and the most preferred shape of the nanoparticulate is an elongated asymmetric shape, such as nanorods or more complex structures involving nanorods, such as nanostars and nanourchins. However, a symmetric composition, such as spherical, particles are not excluded. For example, nanorods with aspect ratio close to one are spheres, which have peak optical absorption in the green spectral range. Formed nanoparticulate clusters may be spherical or aspherical in three-dimensional space, the formed shape may be fractal or chaotic and may be a combination of various aggregate shapes and structures.

The most optimal nanoparticles in terms of maximum absorption in the near-infrared spectral range are the silver nanorods. Also, silver nanorods or more complex elongated silver nanostructures can be designed to absorb near-infrared very strongly, i.e., with an absorption cross-section up to 100 times the physical cross-section. However, silver is not a completely inert metal and can be toxic to normal cells, if used in large concentration. The most optimal nanoparticles in terms of minimal toxicity in the absence of radiation are gold nanoparticles. Furthermore, gold nanorods or more complex structures encompassing gold nanorods, such as gold nanostars or nanourchins, can be designed to absorb near-infrared very strongly such that the absorption cross-section exceeds the physical cross-section several times over, next only to silver nanostructures.

The dimensions of the nanoparticulate are determined from its size which must be sufficiently small to be suspended in a water solution of an appropriate surfactant, e.g., PEG, as compared to pores in biological tissues and blood vessels, so that the nanoparticulate can diffuse through tissue and to be effectively endocytosed by cells, and so that effective targeting of nanoparticles to cell receptors can be accomplished. The size of the nanoparticle must effective for absorbing electromagnetic radiation of chosen wavelength due to plasmon resonance, which requires that the maximum characteristic dimension of the nanoparticulate will be smaller than the wavelength. Some nonmetallic nanoparticles, such as semiconductor carbon nanotubes do not have limitation of size for absorption, but any one skilled in the art of radiation absorption can conclude that there is an optimal size of any particle beyond which absorption of radiation will be less effective and less homogeneous inside the particle. Thus, the nanoparticulate must be no smaller than 1-2 nm and no larger than 1000 nanometers. Nanoparticulate clusters, also provided by the instant invention, may be larger than a single nanoparticulate, i.e. have dimensions of up to several microns.

A preferred nanoparticulate is a partially metallic nanoparticulate with an elongated shape, i.e. with an aspect ratio greater than 1, which may be a collection of nanoparticles. A non-spherical nanoparticulate comprising a nanoparticle aggregate does not require that the nanoparticles of the aggregate be non-spherical. The nanoparticles of the aggregate may comprise spherical nanoparticles ordered in a structure to have the properties of the nanoparticulate disclosed herein. Examples of nanoparticulates are spherical nanoparticles or nanospheres, nanorods or nanoshells.

Particularly, a nanoparticulate aggregate is so ordered and the nanoparticles are at least partially coated with an organic material suitably comprising complementary molecules with high affinity to each other to ordain such order. For example, a collection of spherical nanoparticles may be aggregated as an elongated nanoparticulate, which shift their optical absorption as a function of the aspect ratio, i.e. ratio of small axis length to long axis length. One example of elongated nanoparticulate is gold or silver nanorods. One skilled in the art can appreciate that these types of nanoparticulates have tunable absorption in the near-infrared spectrum of electromagnetic radiation and that their absorption peaks are narrow and very strong, i.e., much stronger than those of biological molecules. These properties are beneficial for application in the Laser Activated Nano-Thermolysis Cell Elimination Technology.

A nanoparticulate used in this invention may be combinations of nanoparticles of one shape with nanoparticles of another shape to form nanoparticulate geometries effective to absorb a selected specific wavelength or range of wavelengths and further form nanoparticulate clusters, which help to produce microbubbles within target cells, which in turn produce maximum thermomechanical damage by laser or other electromagnetic pulses. Thus, for medical or biological applications the details of both dimension and shape are important to LANTCET, since these parameters enable efficient accumulation of nanoparticulate clusters in the target body, such as abnormal biological cell.

The nanoparticulate may be at least partially metallic” and be effective to absorb electromagnetic radiation by the plasmon resonance mechanism. Alternatively, the present invention encompasses nanoparticles, such as carbon nanotubes, that possess properties of semiconductors and yet have very strong optical absorption at various wavelengths of electromagnetic radiation. Either are effective in the LANTCET methods and systems described herein.

Bioconjugation

The present invention encompasses the use of nanoparticles or aggregates of nanoparticles that are conjugated with biological, i.e., organic material. The purpose of such bioconjugation is to (i) produce nanoparticulates that are well suspended in water, (ii) to target specific receptors in abnormal cells and (iii) to form clusters of nanoparticles inside cells or on the cell surface, but not outside the cells in suspension. It is desirable that the nanoparticles have multiple molecules conjugated to their surfaces to optimize the biological and chemical properties of the particles and to maximize the desired formation of clusters inside target cells, but not to form aggregates outside the cells. Coated metal, partially metal and nonmetal nanoparticles or aggregates of these nanoparticles as contrast agents for laser activated nanothermolysis, i.e. selective thermomechanical damage to target abnormal cells, may be used.

Conjugated nanoparticles may have coatings that are covalently bound to the surface of the particles and/or coatings that physically adhere to the surface of the particles. One bond used most frequently in conjugation of gold and other metals to biological molecules is a dative S═ bond provided by the thiol —SH group or sulfhydryl group. U.S. Pat. Nos. 6,821,730, 6,689,338 and 6,315,978 and others (29-34) teach methods of nanoparticle bioconjugation for a variety of biomedical applications. One skilled in the art may predict that numerous techniques exist to conjugate nanoparticles with biological molecules so that such conjugation is chemically stable upon administration of said nanoparticulates in vitro and in vivo. Such conjugated nanoparticulate also must be nontoxic in the absence of radiation and be unrecogniseable as foreign by the human (or animal) immune system to protect them from being scavenged by the immune system before they reach the target body.

In addition proteins or other biological molecules may be used as surfactants. Particularly desirable surfactants are block copolymers, especially block copolymers in which one block is polyethyleneglycol (PEG) (35). PEG can be labeled bi-functionally on opposite sides of the polymer. One side is usually labeled with a thiol or SH group to have strong affinity to metals, especially gold. The opposite side is usually labeled with an NH2 group, which permits convenient conjugation of proteins, such as a monoclonal antibody. PEG, in having hydrophilic groups on the outside of the polymer, prevents nanoparticles from being recognized by neutrophils, macrophages and other scavengers in the circulation or in the mixture of blood cells. The invention further encompasses the use of a nanoparticulate comprising nanoparticles that are stabilized against uptake by the reticuloendothelial system using appropriate surfactants or other particle coatings.

Another purpose of the surfactants or other substances used to coat the particles is to prevent particle aggregation outside the target cells. Aggregation of the individual particles would lead to particle growth and to precipitation of the particles from the suspension that would shorten the shelf life of any conjugate formulation. The instant invention encompasses the use of bioconjugates that are stabilized against particle aggregation and precipitation through the use of surfactants or other particle coatings.

Optionally, the surfactant may serve as a platform for the attachment of other chemical species with desirable biological or chemical properties, which may help to form clusters of nanoparticles inside the target cells. For this purpose, a surfactant or other surface-active agent with reactive functional groups is desirable. As a result, both the surfactant and the attachment site should have reactive functional groups. An optional spacer or linker also should have a pair of reactive functional groups.

An enabling component of nanoparticulate bioconjugates is a targeting vector or moiety. The targeting vectors or moieties may be antibody protein, protein fragments, short peptides or other molecules with a strong or high affinity to target receptors and no or little affinity to target other biological molecules on or in the surrounding medium or body. Regardless of whether the targeting vectors adhere directly to the surface of the nanoparticle or is attached indirectly through the surfactant, the specific receptors for the targeting vectors may be chemical groups, proteins, or other species that are overexpressed by abnormal target tissue or cells. Generally, the receptors may be any chemical or biochemical feature of tissue or cell type to be treated and eliminated. In addition the nanoparticles may be conjugated to a secondary vector or moiety, for example, an antibody or peptide having high and specific affinity to the primary vectors. Such secondary antibody may help to produce aggregates of nanoparticles around the primary targeted nanoparticles on the cell surface.

Specific antibodies for targeting leukemia or other tumor cells for elimination of these tumor cells from the body, such as bone marrow transplants), depend on the type of cells being targeted. Examples include, but are not limited to CD33 and CD123 for acute myeloid leukemia (AML), CD20 for chronic lymphocytic leukemia (CLL), or CD19, CD20 and CD22 for acute lymphoblastic leukemia (ALL).

In addition to successful targeting, clusterization of accumulated nanoparticles in cells of interest must occur. Preferred clusters are two- and three-dimensional structures comprising even numbers of elongated nanoparticles, e.g., two- and three-dimensional stars pyramids or other such structures. Methods for organizing metal nanoparticles into stabilized clusters of aggregates are well-known. For example, covering the surface of different gold particles with complementary strands of DNA favors the self-assembly of the particles into ordered aggregates (36-38). Aggregate formation results from the favorable interaction between the complementary strands of DNA. The instant invention encompasses the use of contrast agents for optoacoustic imaging comprising aggregates of particles coated with complementary strands of artificial or natural DNA, RNA or analogs of RNA or DNA.

Preferably, the nanoparticulate comprises aggregates of metal particles. Such aggregates exhibit a collective plasmon resonance that enhances the total accumulated thermal energy over that expected for single particles in proportion to the total volume of a cluster. Furthermore, clusters of metal nanoparticles can exhibit collective resonance absorption, which is stronger than a simple additive of the optical absorption by single particles (39). For example, a cubic stack of 16 nanorods of 4 layers with 4 nanorods in each layer will absorb more optical radiation than 16 separate nanorods. The presence of a collective plasmon resonance for a collection of nanoparticles is evidenced experimentally by a non-linear increase in the intensity of the optoacoustic signal as the particle concentration increases and aggregates are formed.

A useful biochemical means of promoting controlled particle aggregation is to coat different particles with complementary sequences of nucleic acids, referred to as nucleotides, oligo- and polynucleotides, including deoxyribonucleic acid (DNA) and ribonucleic acid (RNA). Physical means also may be used to promote aggregation, e.g. heating of the nanoparticulates with infrared light (40). Elevated temperature also promotes endocytosis. The endocytosis may be enhanced using an internalizing mab (40).

Stimulated Clusterization

A cluster of strongly absorbing nanoparticles must be formed in order to produce effective local thermomechanical effect, such as generation of an expanding vapor bubble. The cluster has a characteristic dimension, D, comprising small particles with a characteristic dimension, d. A single large particle with a characteristic dimension, D, is not effective as the laser-activated target because of several limitations.

First, large single metal particles do not absorb laser radiation as strongly as a group of smaller nanoparticles because Plasmon resonance interactions are limited to nanoparticles much smaller than the wavelength of electromagnetic radiation. The near-infrared radiation 650-1150 nm penetrates cells with minimum absorption, thus being the most beneficial for selective laser treatment. Nanoparticles that are much smaller than 650-1150 nm are in the range of 10-250 nm, will possess maximum absorption per particle, but are too small to generate vapor bubbles at the temperature of ΔT=100° C., i.e., the temperature of adiabatic vaporization.

A large particle made of organic dyes, polymers or other absorbing materials, can be used to generate vapor bubble at the temperature of 100° C.:


ΔT=μαF/ρC=100° C.  (1),

where μa is the optical absorption coefficient, F is the laser fluence, ρ is the nanoparticle density, and C is the heat capacity of the nanoparticle material. However, due to relatively low absorption of these particles, a very high laser fluence, F, is required to reach 100° C. in these particles. Such high laser fluence is not safe for normal cells and selectivity of the treatment would be lost.

Secondly, large and strongly absorbing particles also are not effective as a contrast agent, because these particles cannot be effectively targeted to cells. The efficiency of targeting is roughly inversely proportional to the particle size. This is explained easily because a large particle cannot be held strongly by a single chemical bond formed between the antibody and the receptor. Also, large particles can not be effectively conjugated to vector molecules, such as monoclonal antibodies or peptides, specific to receptors of abnormal cells. Large particles also are very hard to keep in water suspension with no sedimentation at the bottom of the cuvette. Large particles also will not penetrate through tissues in case their target is not on the very surface.

Clusterization of nanoparticles on and in target cells occurs due to specially designed targeting procedure, which utilizes complementary molecules and high affinity molecular reactions to increase probability of nanoparticle-nanoparticle interaction in cells. The strength of nanoparticle-cell receptor interactions is also maximized by choice of monoclonal antibody. In addition conditions of time duration of targeting, temperature of the body to be targeted, concentration of nanoparticles, and conditions of cellular internalization process or endocytosis must be optimized in order to achieve maximum rate of nanoparticle clusterization. The efficacy of clusterization can be represented as a bell-shaped function of each of these quantitative conditions. One of ordinary skill in the art of cell biology can appreciate that the specific optimal conditions depend on specific medical application.

Various general methods known in the art (33-41) may be used to stimulate clusterization of nanoparticulates in the body. A site of interest, e.g., a tumor cell or tissue comprising the same, may be pretreated with a monoclonal antibody having multiple binding sites is targeted with nanoparticles conjugated to a secondary monoclonal antibody specific for the first monoclonal antibody. Alternatively, nanoparticles conjugated with a primary antibody are targeted to the site of interest followed by targeting nanoparticles conjugated to a second monoclonal antibody to the first monoclonal antibody.

In another alternative method, nanoparticles conjugated to a primary monoclonal antibody and further conjugated to a first aggregating molecule, such as biotin, are targeted to a site of interest. Subsequently, nanoparticles conjugated with the primary monoclonal antibody and further conjugated to a second aggregating molecule, such as streptavidin, are targeted to the avidin-linked nanoparticles. The use of the first and the second aggregating molecules that have high affinity to each other is not limited to biotin-streptavidin linking. One of ordinary skill in the art would be familiar with a variety of complementary chemical or biochemical compounds or compositions that have a very strong affinity to each other, for example, but not limited to, adenine-thymine and guanine-cytosine nucleotides, protein A or immunoglobulin. Furthermore, stimulation of clusterization inside target cells may be accomplished by using an internalizing monoclonal antibody.

A two-stage targeting method, which provides delivery and clusterization of the nanoparticles inside the target cell also is provided. At the first stage, a high concentration of nanoparticles is provided at the outer cell membrane using monoclonal antibodies and specific staining conditions to prevent endocytosis of nanoparticles inside the cell, which is being maintained at a low temperature of 4° C. At the second stage, after washing out unbound nanoparticles from the cell suspension the temperature is raised to 37° C. for an optimal time interval of about 30 min to stimulate the process of endocytosis. The nanoparticles are delivered thereby from the cell outer membrane to inside the cell by endocytosis, including formation of vesicles at the cell membrane. Several nanoparticles in proximity to each other will then be captured at the cell membrane by emerging vesicles, which then deliver nanoparticles inside the cell. As a result, the nanoparticles are concentrated spatially within the vesicles and their spatial distribution inside the cell represents the clusters. Vesicles may exist in the cells for a long time, which to allows further stages of LANTCET to be performed. Also vesicles may deliver nanoparticles to other specific cell compartments where further concentration of nanoparticles may occur.

This protocol can be applied to any type of cell because endocytosis is a universal transport mechanism and vesicles will emerge in any type of cell. The occurrence of the nanoparticulate clusters in the target cells as well as single nanoparticles can be visualized by electron microscopy (44). To avoid nonspecific targeting of normal bodies or microstructures, such as cells), the temperature in the targeting chamber is preferably reduced to 4° C. It is demonstrated herein that minimal or no accumulation of nanoparticles occurred in cells having no specific receptors for targeting vectors used.

Administration of Nanoparticles

The methods and system provided herein are applicable to animal or non-animal bodies, such as cancer cells or bacteria. Thus, without limitation, in terms of medical significance, for example, the body may be an in vivo or in vitro specimen and the object may be a molecule or a virus or bacterium. Alternatively, the body may be an ex vivo specimen, such as a disseminated cancer cell, for example a leukemic cell. The body animate may be an animate human or non human and the object may be biological and comprise a specific tissue, cell or microorganism. For example, the object detected may be a tumor in an animate human or a specific cell, bacteria or virus harmful to the human.

The abnormal body or a cell, which is the subject of LANTCET treatment with the goal of elimination, may be pretreated by specific primary monoclonal antibodies or other vectors that can be further used as receptors for secondary monoclonal antibodies or other vectors in order to form clusters of nanoparticulates in the abnormal body. Those skilled in the art can recognize that a targeting vector against cancer receptor may be used not only for selective delivery of nanoparticles to the target body, but also for direct therapeutic purposes. Such conclusion comes with understanding that targeting vectors, such as monoclonal antibodies, attached to protein receptors on the surface of cancer cells may disable vital functions of those receptors and thereby kill those cancer cells.

An example of such therapeutic action is the monoclonal antibody, trastuzumab, commercially known as HERCEPTIN, raised against receptors associated with HER2/neu gene overexpressed in breast cancer cells and other types of cancer. HERCEPTIN has been successfully used for treatment of metastatic breast cancer (42). As disclosed herein, Laser Activated Nanothermolysis of abnormal cells may be enhanced by pretreatment of the target cells with primary monoclonal antibodies or primary vectors raised against vital receptors on target cells. In association with previously disclosed therapeutic effect of targeting vectors, it is contemplated that the nanoparticulates disclosed in this invention also can be used as an anticancer therapeutic agent. In addition, the designed nanoparticulate can contain an agent molecule to enhance toxicity of such nanoparticulate to tumor cells. Such addition is most desirable if the targeting protocol permits absence of the nanoparticulate accumulation in normal cells.

For imaging a human or a non-human body, many modes of application are possible, depending on the therapeutic requirements. Administration of the therapeutic agent can be systemic or local. Administration can be made intravenously, orally, topically or through direct application of the agent to human or non-human tissue or cells. Local administration of the nanoparticulate agent may be by topical application, by means of a catheter, with a suppository, or by means of an implant or by mixing the nanoparticulate with target bodies (cells) in vitro. Other means of local application will be apparent to those skilled in the art.

Furthermore, the contrast agents may be administered in conjunction with a hyperthermic application, that is, the artificial elevation of the local temperature of an organ or another body part. Hyperthermia accelerates the passage of nanoparticles through the capillaries of the vascular system of growing tumors (43). Hyperthermia also will enhance the uptake of the nanoparticulate by other types of diseased tissue and cells through widening pores and channels in cell membranes.

Any of the many different means of elevating temperature are possible. These include, but are not limited to, the application of thermostatic chambers, the use of focused ultrasound, microwave or RF irradiation. Any or all of these heating procedures can be actively applied while the contrast agent is applied, or, alternatively, the heating can take place up to 24 hours prior to the administration of the agent.

Electromagnetic Radiation

A nanoparticulate is administered to a medium surrounding the body for treatment of the body. Preferably, the nanoparticulate is at least partially metallic, has a formed non-spherical shape having a minimal characteristic dimension in the range from about 1 to about 1000 nanometers and has a formed composition capable of absorbing the electromagnetic radiation and of accumulating thermal energy either in the nanoparticulate or in the body greater than the irradiated body could produce in the absence of the nanoparticulate.

In accordance with the invention, electromagnetic radiation is directed onto the body. The electromagnetic radiation has a specific wavelength or spectrum of wavelengths in the range of about 1 nm to about 1 m which encompasses X-rays to radiofrequency. More preferably, the range is about 300 nm to about 300 mm selected so that the wavelength or wavelength spectrum is longer by a factor of at least 3 than the minimum characteristic dimension of the nanoparticulate. Even more preferably, the spectral range is from green (520 nm) to infrared (1120 nm) wavelengths which can be produced by commercially available lasers. Most preferably, the range of wavelengths is in the near-infrared from about 650 nm to about 900 nm, where tissue and cells have absolutely minimal optical absorption and scattering within which the most preferred gold and silver nanoparticles possess maximum absorption cross-section.

The nanoparticulate made of elongated nanoparticles absorbs the electromagnetic radiation more than would one or more non-aggregated spherically shaped particles of the same total volume with a composition identical to the nanoparticulate. The nanoparticulate by such absorption produces an enhanced thermomechanical effect resulting from the absorption.

The most effective and safe LANTCET procedure can be performed with electromagnetic irradiation of a wavelength that generates microbubbles around nanoparticulate clusters with minimal fluence, i.e., energy per irradiated area, in the targeted body. The irradiation can be generated with a laser, but the invention encompasses the use of any radiation source, regardless of its source. Examples of alternate radiation sources include, but are not limited to, flash lamps, incandescent sources, magnetrons, radioactive substances, or x-ray tubes. The invention encompasses the use in LANTCET of nanoparticulates comprising metal particles or aggregates of metal with electromagnetic irradiation having a wavelength matching the wavelength of peak absorption in the nanoparticulates.

Advantageously, the nanoparticles comprise gold or silver and the wavelength for irradiation is about 520 nanometers to about 1120 nanometers. For example, the wavelength for irradiation is about 520 nanometers to about 1120 nanometers and the nanoparticles in a collection are at least partially gold or silver, are elongated in at least one dimension and have an aspect ratio of at least 2.0. Alternatively, the wavelength for irradiation is about 520 nanometers to about 1120 nanometers, the nanoparticles in a collection are at least partially gold or silver, are elongated, and have a bimodal distribution of aspect ratios. Particularly, one local maximum in the distribution of aspect ratios is about 4 and the other local maximum in the distribution of aspect ratios is about 7. In a multimodal distribution of aspect ratios, the electromagnetic radiation comprises two or more wavelength spreads. In an example for a bimodal distribution of aspect ratios of elongate at least partially gold nanoparticles, one wavelength band is about 690 nanometers to about 800 nanometers and another wavelength band is about 800 nanometers to about 1120 nanometers. Alternatively, the same wavelength range is used and the nanoparticulate is a carbon nanotube, preferably a single wall carbon nanotube.

In a partially metallic nanoparticulate, heating thereof is produced preferably through plasmon derived resonance absorption by conductive electrons in the nanoparticulates. Suitably, the electromagnetic radiation used is pulsed and is emitted from a pulsing laser operating in the near-infrared spectral range. Interaction of nanoparticles with the body being detected produces a shift of the absorption maximum by the nanoparticles for the selected wavelength or spread of wavelengths.

Interaction of Electromagnetic Pulses with Nanoparticles of Various Sizes

Preferably, LANTCET utilizes elongated gold or silver nanoparticles with an absorption peak in the near-infrared, i.e., 620-1120 nm, the most suitable spectral range for deep tissue imaging due to the relative transparency of tissues to near-infrared light. Gold nanoellipsoids or elongated nanoprisms, nanoshells, and other metal nanoparticles absorbing in the NIR can be produced in limited quantities in the laboratory (45-47). Therefore, commercially available spherical nanoparticles as a nanoparticulate are used to demonstrate the feasibility of selective ablation of tumor cells using laser activated nano-thermolysis. The disclosure provided herein will allow one of ordinary skill in the art to predict changes in the results of LANTCET for various nanoparticles based on their optical and thermal properties, shape and dimensions.

FIGS. 1A-1B show the absorption cross-section for gold and silver nanoellipsoids (nanorods) as a function of their diameter employing formulas described in detail elsewhere (39). One of ordinary skill in the art can appreciate that the total optical energy absorbed by a nanoellipsoid increases initially as a cube of the diameter, that is, proportional to the volume of these nanoparticles, and then increases as a square of the diameter, that is, proportional to the area of the nanoparticles.

Superheating of the nanoparticle, which evaporates a layer of surrounding water, may produce a microbubble. Surprisingly, it was discovered herein that microbubbles cannot be produced around small nanoparticles, such as about 10 nm to about 100 nm, using optical fluence that heats these nanoparticles up to 100° C., the boiling point of water, and even up to 374° C., the critical temperature of water. Even with extremely high fluence of pulsed electromagnetic radiation, only nanobubbles invisible by optical microscopy can be produced. With a further increase of absorbed energy, the small nanoparticles evaporate and disappear from aqueous suspension. Even for larger nanoparticles (>100 nm) it is statistically difficult to produce visible microbubbles.

The reason for such phenomenon is that a nanoparticle with small volume can accumulate only a limited amount of thermal energy, not sufficient to evaporate a volume of water required to produce a microbubble that can sustain in suspension for a measurable time. There are two major physical reasons for such effect. First, strong surface tension that is inversely proportional to the bubble radius makes the surface tension force very strong for bubbles smaller than certain radius. Secondly, viscosity of water provides an extremely strong force on small nanobubbles thereby preventing their growth to microbubbles. Thus, only when the nanoparticulate cluster has sufficient size and mass to accumulate the thermal energy from electromagnetic radiation, is it possible to generate microbubbles of vapor using energy fluence that heats the nanoparticulate cluster to a temperature between 100° C. and 374° C. This energy fluence under optimal experimental conditions can be safe for normal cells. In the absence of nanoparticle clusters, the required energy fluence will be much higher than 1 J/cm2, which is not safe for normal cells.

Using the known absolute values for gold nanoparticle absorbance along with the thermal diffusion models for different shapes of heated objects (48), an estimate can be made of the minimum laser fluence required to heat a nanoparticle or a nanoparticulate cluster up to a boiling point of water at 100° C. or a critical point of water around 374° C. Upon irradiation with a laser pulse, heat diffusion within a gold nanoparticle occurs on the scale of picoseconds. Therefore, gold nanoparticles will be homogeneously heated with a 5 ns to 50 ns pulse width of a typical Q-switched laser suitable for LANTCET.

Heat diffusion time is the time required for the transfer of about ⅔ of the thermal energy stored in a nanoparticle to the surrounding medium. The heat diffusion time from a nanoparticle or a cluster of nanoparticles to the surrounding water occurs on the scale of sub-nanoseconds to tens of nanoseconds, depending on the size of a nanoparticle, d, being heated by radiation, and its shape (48):

τ HD = d 2 24 χ , if the shape is spherical ( eq . 2 a ) τ HD = d 2 16 χ , if the shape is cylindrical ( eq . 2 b ) τ HD = d 2 4 χ , if the nanoparticulate cluster is shaped as disk ( eq . 2 c )

In the formulas (2a-2b-2c), χ=1.3·10−3 cm2/s is the thermal diffusivity of water at room temperature. The expression describing the increase of temperature of the nanoparticle during laser pulse, i.e. when the particle simultaneously absorbs light and diffuses heat) can be presented in the following fashion (13):

Δ T NP = F σ a NP V NP ρ NP C NP × ( τ HD τ L ) × [ 1 - exp ( - τ L τ HD ) ] ( eq . 3 )

where F[mJ/cm2] is the incident (upon the nanoparticle) laser fluence and σαNP is the plasmon-derived absorption by the spherical gold nanoparticle at the wavelength of laser irradiation, VNP is the volume of nanoparticle being irradiated, ρNP is the density of nanoparticle material (for gold ρg=19.3 g/ml), CNP is the heat capacity of nanoparticle material (for gold Cg=0.128 J/g0K), τHD is the effective heat diffusion time from gold into water (surrounding medium). As formula (3) indicates, for effective utilization of the electromagnetic pulse (laser pulse) energy, the pulse duration must be shorter than the heat diffusion time, τHD. A typical heat diffusion time for preferred nanoparticles with dimensions of 10-100 nm is in the range of nanoseconds. Therefore, an electromagnetic pulse of near-infrared radiation from a q-switched laser with a duration of a 3-10 ns may be an example of a preferred pulse duration.

Sometimes, however, for the sake of cost reduction, pulsed optical sources are being replaced with continuous wave sources (11,16). In case of significant contrast between tumor cells and normal cells, one of ordinary skill in the art can design a successful treatment procedure even using continuous wave (over 1 sec long pulses). Nevertheless, short pulses having duration equal or shorter than the time of thermal diffusion from the heat source to surrounding tissue, will result in much more effective thermomechanical interaction and much better spatial confinement of the damage effects (2-4).

Formula (eq. 3) is true until the temperature reaches 100° C. Then, any additional absorbed energy may contribute to the evaporation of the water around the particle, as well as heating the particle above 100° C. Employing ΔT=80° K, the difference between room temperature and boiling temperature of water in formula (eq. 1), minimal optical fluence required for generation of vapor bubble around absorbing nanoparticles can be calculated.

For example, assume the diameter of a spherical gold nanoparticle is d=200 nm. The heat diffusion time from this particle to water equals 12.8 ns. Then for a typical laser pulse of 12.8 ns in duration, formula (eq. 3) will yield F=0.6 mJ/cm2, which is the critical fluence needed to heat up the spherical gold nanoparticles to 100° C. It can be predicted that any fluence larger than 0.6 mJ/cm2, will result in superheating and possibly evaporation of water around the nanoparticle. Assuming that evaporation, at least in the stage of a thin nanobubble around nanoparticle, does not prevent the nanoparticle from being further heated, the optical fluence that corresponds to a temperature increase in a 200-nm diameter nanoparticle to 374° C. equals 2.6 mJ/cm2. This is a temperature at which conversion of water into vapor occurs instantly.

Surprisingly, however, no bubbles were detected at this level of fluence of laser irradiation. FIG. 3 depicts a magnitude of optoacoustic signal as a function of laser fluence. No deviation occurs at 2.6 mJ/cm2 from a typical linear curve describing the thermoelastic expansion of water. This means that at this level of optical fluence microbubbles do not occur. If any nanobubbles occur, they cannot contribute to the optoacoustic signal or to thermomechanical damage to surrounding the body. σα=1.4×10−9 cm2. A sharp increase in the signal occurs at about 1 J/cm2 indicating contribution of vapor bubbles to the optoacoustic signal. Note that the threshold of deviation of the optoacoustic signal from the linear curve of thermoelastic expansion occurs at a fluence much greater than the fluence, F=0.0026 J/cm2, corresponding to the critical temperature of water, 374° C.

Results presented in FIG. 3 and the calculations above show that the threshold of microbubble generation should decrease with increase of the size of nanoparticles. On the other hand, a very strong absorption of electromagnetic radiation due to plasmon resonance may be lost by a large nanoparticle, e.g., microparticle, since plasmon resonance theory requires that the size of the nanoparticles must be smaller than the electromagnetic wavelength. Furthermore, microparticles are not practical from the clinical prospective, since these particles are too large (heavy) to be suspended in water and can not propagate through biological cells and tissues, which makes targeting of these types of particles to a cell or other biological body difficult or impossible. Based on these considerations, clusters of nanoparticles can be used to effectively target cells and then to effectively generate vapor microbubbles by a low fluence laser radiation. If an ideal size cluster can be formed selectively in target cells, the laser fluence required for cell damage will be in the range of only a few mJ/cm2, which is absolutely safe for normal cells and tissues.

FIG. 4 demonstrates the threshold of microbubble formation from laser irradiated clusters of gold nanoparticles. A targeting protocol is designed, which resulted in the accumulation of clusters of about ten 30-nm diameter nanoparticles in the target tumor cells, human B-lymphoblasts. The threshold fluence of the cell damage, confirmed with observation of the microbubbles by photothermal detection, was found to be about 100 mJ/cm2, which is 50-60 times lower than that observed for individual gold nanoparticles with diameter 200 nm and 30 times lower than the cell damage threshold for cells nonspecifically targeted with the same 30-nm diameter nanoparticles. Both the microbubble generation threshold and the threshold of cell damage was always significantly higher in cases when no clusters in cells was observed.

FIGS. 1A-1B showing absorption coefficients of gold and silver nanorods and expressions (eq. 1), (eq. 2b) and (eq. 3) permit estimation of a nanoparticle temperature for a nanoellipsoid of revolution. Assume a gold nanorod with diameter of 20-nm and length of 100-nm. Its absorption cross-section is 8.5×10−10 cm2, i.e. almost equal to that of a gold nanosphere with diameter of 200 nm, while its volume is only 200 times smaller than that of the nanosphere. Significantly stronger optical absorption of gold nanorods compared with gold nanospheres of equal volume results in a dramatically reduced minimum laser fluence required to heat up nanorods to a temperature required for vapor bubble formation. Thus, clusters of elongated metal nanoparticles are preferred over clusters of spherical metal nanoparticles for producing near-infrared radiation induced microbubbles. FIGS. 2A-2C show that experimentally measured optical properties of actual gold nanorods are very close to those calculated theoretically. A controlled fabrication of relatively monodisperse nanorods with an aspect ratio not deviating from the maximum can be achieved. Furthermore, conjugation with PEG slightly changes spectral width and position of the peak absorption.

Recent developments in nanotechnology allow for engineering nanoparticles with various shape and dimensions thereby facilitating the development of new imaging and therapeutic nanoparticulates. Colloidal gold is especially attractive since it is inert material that has been used for therapeutic applications (49-50). It is noteworthy to mention that upon intravenous injection of gold nanoparticles, unattached nanoparticles that could be a source of background noise for LANTCET can be rapidly removed from the circulating blood pool by liver and other organs of the reticuloendothelial system (51).

In summary total absorbed energy of electromagnetic radiation is proportional to the volume of a nanoparticulate. Thus, the total thermal energy stored in said nanoparticulate is also proportional to its volume. Heat diffusion rate decreases with nanoparticle dimension to the second power. Probability of microbubble generation is inversely proportional to dimension, i.e., surface tension, and viscosity that strongly affects the bubble generation is proportional to the nanoparticle dimension to the second power. These factors require that a nanoparticulate cluster is formed to reduce the threshold of microbubble formation, that can be used for LANTCET.

Real-Time Imaging and Monitoring

Lapotko et al (21-23) teach a method and apparatus for obtaining an image of a body, e.g. a cell, to allow examination of a number of submicron heterogeneities simultaneously, such as microbubbles or heated areas, including a method for detecting size. A relatively large sample surface, bigger than the cell diameter, is exposed to the pump laser radiation. The size of the surface exceeds the wavelength of the pump laser beam used. In fact, a surface of any size could be irradiated, but, logically, the size could not exceed the size of the sample itself because the chosen probe laser beam diameter is not smaller nor comparable with the pump laser beam diameter and is not larger than the maximum overall dimensions of the sample.

Spatial distribution of absorbing heterogeneities, e.g. bubbles, in the irradiation zone is determined by the synchronous measurement of the diffraction limited phase distribution through the whole cross-section of the probe laser beam which is transformed into an amplitude image by a phase contrast method. The size of separate microheterogeneities larger than the pump laser beam wavelength is determined by analyzing the amplitude image structure. The amplitude image corresponds to the refraction index change distribution induced by the pump laser in the object observed.

The average size of microheterogeneities smaller than the wavelength is measured indirectly by the characteristic time of cooling which is dependent on the size. That measurement is based on the speed measurement of the phase change of diffraction-limited images of those microheterogeneities at various points of the probe laser beam cross-section at different moments in time. Measurement begins right after pump laser irradiation has taken place as the chosen irradiation period is much shorter than the characteristic time of cooling of the microheterogeneity observed.

A short-time irradiation can be performed by two methods. The first method uses a single laser pulse. It is the duration of the pulse that determines the period of effect. A pulse-periodic mode, usually with a porosity greater than 1, also could provide this effect. The second method uses continuous laser pump radiation that is intensity-modulated with a relatively high modulation frequency ranged from a few kHz to hundreds of MHz. In this case duration of a single effect is determined by a modulation semi-period. This effect repeats with the frequency determined by the laser modulation frequency. Information about the time of cooling would be carried by the probe laser beam time phase related to the pump laser beam time phase.

A number of versions of the probe beam realization could be used. For example, a part of the pump laser beam could be used as a probe beam. Propagating the probe beam through an additional optical delay line regulates its delay time as related to the main beam. The chosen probe beam intensity should be considerably (at least 5-10 times) lower than the main beam intensity, so as to have minimal effect on the measurement results.

The phase distribution of the probe laser beam in the function of the pump laser wavelength should be measured, so as to obtain information about spectral properties of separate microheterogeneities simultaneously with their sizes. It is suggested that said measurement is to be accomplished by at least two time delays as related to the pump beam pulse at every pump laser wavelength. Dynamic change of the microheterogeneities or microbubbles, induced by the pump beam, is examined by changing at least two phase images, one of which is obtained immediately before the pump pulse operation and the other obtained simultaneously with the pulse operation or with a delay, with their subsequent subtraction. This is particularly important in case there are insignificant alterations of the image structure which are difficult to identify using only one image, as the measurement precision is low.

For the LANTCET method, an additional optical system of phase contrast is used as an optical transformation unit to transform phase distribution in the probe beam cross-section to an amplitude image. A registration unit is a high-speed multi-channel photodetector, for example CCD-matrix, in the pulse mode to register the amplitude image of the probe beam at various moments of time as related to the moment of the pump laser pulse operation. Another version of the registration unit is a number of one-channel photodetectors used to register time amplitude changes for one or several zones in the amplitude image of the probe beam. The probe beam falls simultaneously on all the detectors due to a semi-transparent system of mirrors placed in the way of probe beam behind the phase-contrast system. Another solution is a consecutive spatial shift of said detector using an additional switch unit.

For examination of the microheterogeneities considerably smaller than the wavelength, the fundamental solution is to introduce a synchronizing unit and a time delay regulating unit connected with each other, with the pump laser units, with the probe laser forming unit, and with the registration unit. A gradually regulated delay provides, when using a probe laser and pump laser pulse regimens, a precise measurement of the cooling time for the absorbing heterogeneities heated by the pump pulse to estimate the average size of the heterogeneities. If the continuous mode of the probe laser is used, a synchronizing unit that switches the probe beam phase monitoring at the moment of the pump laser pulse operation is used.

In the continuous mode of the pump laser with intensity modulation, the device contains an additional intensity-modulating unit placed in the path of the pump beam distribution. Registration of the continuous probe beam modulation caused by pump radiation, via refraction index modulation, is provided by the synchronous integrating unit connected to a photodetector or multi-channel photodetectors of the probe beam. The unit also receives a signal from the pump beam modulator. The necessary information about the pump laser (time) phase is carried by the signal.

Also provided is a one-channel mode, where pump radiation accomplishes the probe beam functions simultaneously. In this case phase distribution in the pump beam cross-section itself is registered. The filter cutting the pump beam in front of the photodetector should be removed to follow this scheme. The system of splitting the pump laser beam into a main beam and an additional beam is introduced into the path of the pump beam, coming from the pump laser to use part of the pump laser as a probe beam, the additional beam accomplishing the probe beam function. The optical delay line connected with the time delay unit is introduced in the path of the probe beam.

A probe beam-forming unit can be realized both as a continuous laser connected with the synchronizing unit and as a pulse laser connected with a time delay unit. A probe beam turning unit should be introduced as related to the sample observed to obtain a three-dimensional tomographic image. Another version is to introduce the turning unit of the sample itself, where the turning unit is connected with the synchronizing unit. The device can also be equipped with an image processing unit connected with the photodetectors, the synchronizing unit, and the time delay unit. Its functions include image comparison at various moments of time, and another one is comparison of photothermal and regular optical images.

The device is additionally equipped with a pump beam wavelength-changing unit connected with the pump laser unit. Various methods can be used to provide the laser wavelength change. These methods include temperature and pressure influence on the active element; using spectral elements within a pump laser resonator in the form of a prism, diffraction grid, interference filters, or other applicable elements.

LANTCET begins when the object containing absorbing microheterogeneities is irradiated with a probe laser beam where the chosen probe beam diameter is not smaller than the pump beam diameter and is not larger than the maximum overall dimensions of the sample. Intensity of the probe beam should be considerably, i.e., at least 5-10 times, smaller than the pump beam intensity to cause minimal effect on the measurement results. The diffraction-limited distribution of the probe laser beam phase over the whole cross-section is then transformed to an amplitude image using the phase contrast method. The obtained values of the probe beam phase φ0(x, y) and the phase-corresponding amplitude I0(x′, y′) of the image are basic for further analysis.

The next step is irradiation of the object containing absorbing microheterogeneities by a focused pump laser beam having a short pulse width and a wavelength coinciding with the absorption line of the microheterogeneities. The pump pulse immediately irradiates a relatively large sample surface, where the size of the surface is larger than the wavelength of the pump laser used. In fact, the surface could be of any size, but logically, it couldn't be larger than the sample itself. If such an effect occurs, light energy absorption in the sample is not uniform: microheterogeneities absorb light most actively. Thus, a live cell has various absorbing structures, e.g., cytochromes, organelles, and mitochondria, the size of which vary from a few nm to hundreds of nm, i.e., considerably smaller than the average size of a cell (5-20 microns). However, their ability to absorb light is so high, that it causes thermal effects causing a temperature rise 10-1000 times higher than the temperature of the cell's environment. Cooling of the structure that has absorbed light energy begins by heat diffusion after the end of the pump pulse operation.

The time of cooling for a single sphere-looking object is:

τ T = R 2 6.75 K , ( eq . 4 )

where

tT is the time of cooling of the object, sec;

R is the radius of the object, m; and

K is the temperature conductivity coefficient (m2/c).

This time equals 10−5-10−4 sec for the majority of blood cells.

A primary thermal response could be presented as the distribution of temperatures over the x-axis and the y-axis:

Δ T ( x · y ) = α ( x · y · λ ) ɛ ( x · y ) ρ C , ( eq . 5 )

where

ΔT is the distribution of temperatures over the x- and y-axes,

α is the light energy absorption coefficient with the wavelength,

ρ is density, kg/m,

ε is energy density in the pump beam, J/m2,

C is thermal capacity, J/kg° C.

Expressions (eq. 4) and (eq. 5) help estimate the temperature effect both for an object (a cell) as a whole and for its structural elements, i.e., absorbing heterogeneities. The intensity of the effect depends on a specific absorption coefficient and the heterogeneity's size. tT would be about 10−8 s and less for the submicron structures having the sizes smaller than the wavelength (10−7−8 m). It means that significant rise of the local temperature could be achieved only providing the pump pulse width or the modulation period T=1/f (where f is a modulation frequency) are smaller or at least comparable with tT. Otherwise the local temperature effect, being the very source of the local heat variations of the refraction index, would not be achieved. The local heat variations of the refraction index Δn(x, y) induced by the pump pulse absorption could be presented as follows:

Δ n ( x , y ) = α ( x , y , λ ) ɛ ( x , y ) ρ C ( n T ) p , ( eq . 6 )

where

α is the light energy absorption coefficient with the wavelength,

ρ is density, kg/m,

ε is energy density in the pump beam, J/m2,

C is thermal capacity, J/kg° C.; and

T is temperature.

A further step comprises irradiating the object containing heated absorbing microheterogeneities by the probe laser beam where the chosen probe beam diameter is not smaller than the pump beam diameter and not larger than the maximum overall dimensions of the sample. Intensity of the probe beam should be considerably, i.e., at least 5-10 times, smaller than the pump beam intensity to cause minimal effect on the measurement results. The phase of the probe beam wave front will be distorted from the local heat variations of the refraction index when the probe beam would propagate through the sample. The phase deviations φΔ(x, y) could be described as follows:

ϕΔ ( x , y ) = L α ( x , y , λ ) 2 π ɛ ( x , y ) λ 0 ρ C ( n T ) p , ( eq . 7 )

where

L is the geometrical length of the probe beam way in a heterogeneity,

n is the refraction index,

Δn is the refraction index variations on the x- and y-axes,

α is the light energy absorption coefficient with the wavelength,

ρ is the density, kg/m,

ε is the energy density in the pump beam, J/m2,

C is thermal capacity, J/kg° C., and

T is the temperature.

In a further step the diffraction limited phase distribution of the probe laser beam over the whole cross-section is transformed to an amplitude image using the phase contrast method. Taking into account the values φ0(x,y) and I0(x′,y′) previously obtained in an unexcited state, parameters of the probe beam propagated through the exited sample at the moment of time t0 could be presented as follows:


φ(x,y)=φ0(x,y)+Δφ(x,y)  (eq. 8),

where

φ0(x, y) is the probe beam phase in the absorbing zone of an unexcited object

Δφ(x, y) is alteration of the probe beam in the absorbing zone of an unexcited body,


I(x′,y′)=I0(x′,y′)+S(x′,y′)  (eq. 9),

where

I0(x′, y′) is the amplitude of the photothermal signal in an unexcited state,

S (x′, y′) is the required photothermal signal being subject to registration and analysis.

The size of separate microheterogeneities larger than the pump laser wavelength is determined using structural analysis of the amplitude image measured immediately at the moment of the pump laser operation and corresponding to the refraction index change distribution induced by the pump laser in the object observed. To determine the average size of the microheterogeneities smaller than the wavelength, the phase alteration speed of the diffraction-limited images of said microheterogeneities in various points of the probe beam cross-section should be measured. Measuring begins immediately after the pump pulse effect has taken place. To achieve this, the probe beam irradiation and the phase distortion analysis of the object are performed a number of times, e.g., at two moments of time.

The theoretical limit of this method is conditioned by the terminal time of transformation of optical energy to thermal energy which is 10−13 seconds for condensed mediums. This time corresponds to 1 Å and could be achieved using femtosecond lasers. The photothermal signal amplitude S is proportional to the temperature change in the absorbing zone and decreases due to thermal conductivity.

U.S. Pat. No. 5,840,023 teaches the acquisition of optoacoustic images with contrast agents. In this method, a short pulse of irradiation is followed by detection of the induced pressure wave, which is then used for generation of an image. In the practice of the present invention, the pulses of electromagnetic radiation preferably have a duration of about 10 ns to about 1000 ns. When the tissue to be imaged is simulated by a solid slab tissue, the radiation fluence on the surface of the slab is about 10 mJ/cm2. For other configurations of the test sample or for living human or non-human bodies, the surface fluence will vary, but will always be in the range of about 1 to about 100 mJ/cm2, which generally is considered safe.

In addition both photothermal and optoacoustic imaging, either simultaneously or sequentially, can be utilized to monitor pulsed laser interactions with nanoparticulates and their clusters in the course of LANTCET procedure. The monitoring method may include photothermal imaging and microscopy, optical and optoacoustic detection of thermal field, thermal lens and bubbles generated around nanoparticulates. The imaging can occur during administration of the nanoparticulate, immediately after administration, or at some later time to allow for accumulation of the agent in the target cell or other body.

LANTCET System

FIG. 5 shows a designed system that utilizes the LANTCET method of laser activated nanothermolysis for purging tumor cells. For example such system may be a modification of a blood dialysis system where cells containing or accumulating bioconjugated nanoparticles can be irradiated extracorporally, i.e. outside human or animal body. The main components of the system 10 are the cell chamber 1, a source of targeting moieties 2, e.g., monoclonal antibodies or peptides, and nanoparticulates 3, an optical chamber 4, such as an optical flow cuvette, a pulsed source of electromagnetic radiation 5, and a means for filtering 6 and collecting 7 the products of damaging thermomechanical effects, e.g., a hemosorption system or similar system. The cell chamber 1 is independently fluidly connected to the source of targeting nanoparticulates 2,3, to the optical chamber 4 and to the filtering and collecting means 6,7. The pulsed source of electromagnetic radiation 5 is positioned to irradiate the cells in the optical chamber 4. Optionally, an imager 8 is positioned to receive a photothermal signal or for generating an optical image of thermomechanical effects affecting the cells in the optical chamber 4.

The system is configured so that bioconjugated nanoparticles can selectively target the cells in the cell chamber whereby nanoparticulate clusters accumulate in the cells. The targeted cells containing the nanoparticulate clusters flow to the optical cuvette where they are irradiated with laser or other electromagnetic pulses. The laser pulse heats the nanoparticulates which heat subsequently induces microbubble formation. The formation of microbubbles causes selective and increased levels of thermomechanical damage. Optionally, the imager may monitor the process of cell nanothermolysis by laser-induced microbubbles and/or direct the damaging thermomechanical effect against the targeted cells.

The following examples are given for the purpose of illustrating various embodiments of the invention and are not meant to limit the present invention in any fashion.

Example 1 In Vitro Laser Activated Nanothermolysis of Tumor Cells Using Nanospheres

An in vitro model with myeloid K562 cell line, cryopreserved human cells, tumor cells (patient-derived acute B lymphoblast leukemia), both referred to as tumor cells, and normal stem cells are used. Well-defined specific MABs are used as primary MAB for targeting, i.e., CD15 and Glycophorin-A for K562 cells and CD19 for acute B-lymphoblast leukemia cells. Selection of MABs was realized by using flow cytometry. MABs CD 15 and Glycophorin A were selected for K562 cells based on their superior level of expression on surface of those cells. CD 19 was selected in the same way for acute B lymphoblast leukemia (ALL) cells. For controls for the single-cell model, untreated K562 cells are used and for the suspension model, normal stem cells are used.

Specific Targeting Protocol

Cells at a concentration of 800,000/ml in phosphate buffer solution (PBS) with 1% fetal bovine serum (FBS) were preincubated with two primary MABs CD15 and Glycophorin-A (20 ml/ml of each) (K562 cells in single-cell model) and CD19 (acute B lymphoblast leukemia cells and normal stem cells in suspension model) for 30 min with shaking in the dark at 4° C. After incubation, the cells were centrifuged at 300 g for 4 min and washed twice. Then the cells were incubated again for 30 min with 2×1010/ml of 30 nm gold NPs conjugated to secondary MAB, Goat anti-Mouse IgG (British Biocell International, Cardiff, UK via Ted Pella Inc., Redding, Calif.). Cells were separated from free uncoupled NPs by centrifugation at 300 g for 4 min. The bottom pellet of cells was resuspended in PBS-FBS solution and immediately used. The same cells were used for preparation of cell samples with modified protocols.

Non-Specific Targeting Protocol

The same procedure as described above was applied, however, without primary MAB, only with secondary MAB-NP complex. As a simple control, bare gold NPs without any MAB were incubated with cells under the same conditions. To determine cell viability, trypan blue dye was used according to a routine protocol, on an aliquot of cells from the stock suspension before the laser treatment and on every treated sample after the laser treatment. Four K562 cell samples were prepared as for the single-cell model. Two cell samples (leukemic and normal stem cells) were prepared for the suspension model using the specific targeting protocol. Additional control sample of nanoparticles in solution without cells was also prepared.

Laser Treatment Procedure and Cell Damage Measurements

LANTCET was experimentally studied in the two models, single cell and suspension of cells. FIGS. 6A-6B depict a schematic diagram of LANTCET with simultaneous monitoring and guidance by photothermal microscopy. In the diagram a mixture of cells is stained with MAB-conjugated gold nanoparticles that selectively attach to target receptors and form clusters (FIG. 6A, right), cells are irradiated by a laser pulse that induces the bubbles around clusters of nanoparticles (FIG. 6B), the bubbles destroy only cells targeted with nanoparticles and non-targeted cells (FIG. 6C, left) remain intact; and a photothermal image shows tumor cells loaded with nanoparticles prior to pulsed laser irradiation and fragments of cells destroyed by the laser pulse (FIG. 6D). In the first model, single cells were irradiated one by one and any possible damage was detected immediately. After the incubation with gold nanoparticles, the samples of K562 cells were immediately placed in the sample chamber (S-24737, Molecular Probes, OR) mounted on a microscopic slide to produce a monolayer of cells. Individual cells (total of 150) were irradiated within 7 min with 10-ns long single focused laser pulse at 532 nm. The green light at 532-nm is strongly absorbed by nanoparticles, but not by the cells. All samples were exposed to the same laser irradiation conditions, i.e. a specific optical fluence.

A photothermal (PT) microscope was used for visualization of vapor bubbles around the laser irradiated NP [30]. Photothermal microscopy may be used not only for detection of microbubbles around the nanoparticles, but also for the real-time monitoring of the laser-induced damage of individual cells with nanosecond temporal resolution. Single cells were irradiated with two collinear focused laser beams, i.e., a pump pulsed beam for bubble generation and a probe low intensity continuous-wave beam for detecting transient thermomechanical changes in the irradiated cell. Any change in the optical refraction index caused by a thermal field or a bubble produces a thermal lens effect that influences the probe-beam intensity at the input of a photodetector. Any bubble-specific photothermal response signal detected from a cell at the same time indicates damage of the irradiated cell [22].

Cells were irradiated one by one with single focused laser pulse at 532 nm (10-ns duration) and the photothermal response from each cell was recorded with a photodetector measuring changes in the power of the probe laser at 633-nm (temporal profile of probe laser signal) [21,22]. The probability PRB of bubble generation at a specific laser fluence was measured as:


PRB=N2/N  (eq. 10)

where N is the amount of irradiated cells and N2 is the amount of cells with bubble-specific photothermal responses. The bubble generation threshold fluence for a specific pump laser wavelength was defined as the fluence that corresponds to a PRB level of 0.5.

In the second model, feasibility of LANTCET for purging was evaluated using suspension of cells. Suspensions of test and control cells were injected into rectangular glass chambers and then were irradiated with single laser pulses. A laser beam with diameter of 1 mm was scanned along a cuvette with dimensions of 0.4×0.1×10.0 mm for about one minute for each sample. After the irradiation, the laser-induced cell damage was assessed within 5 min in the same cuvette using the standard trypan blue dye, optical microscope and digital camera for visualization and counting the level of positively stained cells. The damage probability, DP, was then calculated.

Electron Microscopy of Leukemic Cells

Binding of IgG-conjugated gold nanoparticles to primary MABs attached to CD15 and Glycophorin-A receptors on the surface of K562 cells causes internalization of the complex. Whether phagocytosis took the NPs into the cells was examined using transmission electron microscopy. The targeted cells were fixed in situ in a mixture of 2.5% glutaraldehyde and 1.5% formaldehyde in 0.05 M cacodylate buffer pH 7.2 to which 0.03% trinitrophenol and 0.03% CaCl2. After washing in 0.1 M cacodylate buffer cells were pelleted and processed further as a pellet. They were post-fixed in 1% OsO4 in 0.1 M cacodylate buffer, en bloc stained with 2% aqueous uranyl acetate, dehydrated in ethanol and embedded in Poly/Bed 812 (Polysciences, Warrington, Pa.). Ultrathin sections were cut on Reichert-Leica Ultracut-S ultramicrotome, stained with lead citrate and examined by a transmission electron microscope JEM 100 CX II (JEOL, Japan). Amount of nanoparticles as per cell was estimated through counting particles in four Emimages each visualizing a thin (60 nm) slice of a cell, averaging counts in 10 different cells (total of 40 microscopic slides were processed) followed by extrapolation of obtained numbers to the total volume of the cell.

Evaluation of Add Nanoparticle Uptake by Cells with Electron Microscopy

Electron microscopy (EM) images of K562 cells (FIGS. 7A-7B) revealed that nanoparticles fill the entire volume of the cell and do not concentrate only on its outer membrane where target receptors are located. Clusters of closely packed 5-20 particles were found in cells that were selectively targeted using the primary MAB and bioconjugates of nanoparticles with the secondary monoclonal antibodies (FIG. 7B). The total average number of nanoparticles per cell estimated by extrapolating the number of nanoparticles counted in 60-nm thin slices corresponding to EM images was about 31650±4810 in specifically targeted cells. In contrast, a much smaller number of nanoparticles was found in non-specifically targeted cells, i.e., about 1500±350 nanoparticles per cell and only in about 200±90 nanoparticles per cell in control cells incubated with bare nanoparticles. No clusters were observed in non-specifically targeted and control sample (FIG. 7A). High nanoparticle contrast (ratio of NP level for specific targeting to that of non-specific targeting) was 21. This ratio further increased to 158 for specific targeting relative to the control of bare nanoparticles. The selectivity achieved by specific targeting of tumor cells indicates a high degree of LANTCET safety in that only cells with clusters of nanoparticle may be destroyed while leaving other cells not damaged by laser pulse.

Laser Elimination of Tumor Cells-Single Cell Model

Using the single cell model, PT responses from individual K562 cells to pulsed laser radiation (532-nm, 10-ns) were studied. PT responses from the free space between the cells (background), from individual cells targeted with nanoparticles and from suspension of single nanoparticles without cells were obtained. The microbubble generation and cell damage were monitored simultaneously with the probe laser pulse by analyzing profiles and amplitudes of PT responses as depicted in FIGS. 8A-8C. Among all studied samples, only the cells selectively targeted with gold spherical nanoparticles showed prominent bubble-specific PT response (FIG. 8C). These cells also frequently exhibited apparent visual signs of the damage, such as fragmentation after irradiation with a single laser pulse. Microbubbles with a duration 1-3.5 μs (FIG. 8C) were detected only in specifically targeted cells, while in other test samples such long microbubbles were never observed even at optical fluences 10 times exceeding the threshold of bubble generation. Long bubbles were detected only among those cells that manifested the clusters of nanoparticle in their EM images.

The explanation for this observation is found in the previously discussed models of vapor bubbles generated around superheated nanoparticles. The temperature of a laser-heated nanoparticle is proportional to the ratio τHD/τL of heat diffusion time and the laser pulse duration, where the heat diffusion time from a spherical nanoparticle into aqueous medium can be found as the ratio of nanoparticle diameter squared and the heat diffusion coefficient (χ=0.0013 cm2/s) in water τHD=d2/24χ[54]. Heat diffusion time equals 0.3 ns for a single gold nanoparticle with diameter of 30-nm and this time is much less than laser pulse length (10 ns). Such rapid heat diffusion distributes the laser pulse energy over a larger volume, so that only a 0.033 fraction of laser energy is used to heat the nanoparticle. More importantly, for generation of vapor bubbles around nanoparticles, the surface tension and the dynamic viscosity of water prevent formation of expanding bubbles around nanoparticles superheated substantially above the critical point temperature of water (647° K).

The situation changes for NP clusters with effective diameter of more than 300 nm consisting of about 10-20 nanoparticles, so that the initial surface tension is 10 times lower and the full thermal energy of a laser pulse is utilized to heat the cluster of gold nanoparticles. Therefore, the threshold of pulsed laser interaction with clusters of nanoparticles is significantly lower than that for a single nanoparticle. Superheating of the nanoparticle clusters generates a much bigger vapor microbubble capable of damaging even large cells. Nanoparticle clusters were destroyed by single laser pulses with optical fluence of 5 J/cm2, which could superheat NP to a temperature significantly above the boiling temperature of gold (287° K). Laser pulses that followed the first microbubble-generating pulse produced no bubbles, and no bubble-specific PT response was detected for the second pulse (FIG. 8A).

It is important to note the difference in PT response amplitude and lifetime, the parameters dependent on the bubble diameter, between the control suspension of nanoparticles (FIG. 8B) and the cells specifically targeted with gold NPs (FIG. 8C). Small amplitude and a short 0.3 ms duration of the PT response signal detected in the suspension of gold nanoparticles (FIG. 8B) indicates that only a very short-lived nanobubble was generated even with optical fluence of 35 J/cm2. In contrast, cells specifically targeted with nanoparticles (test sample) produced a much stronger PT response (FIG. 8C), which indicates that the bubble was almost one order of magnitude larger and, therefore, lasted longer. Such increase in the bubble lifetime and size may be explained only by formation of clusters of NPs in tumor cells.

Cell damage probability DP for test and control samples was analyzed at several different levels of laser fluence: 5, 35 and 90 J/cm2 (FIG. 9). Both the DP level and the laser fluence threshold for bubble generation depend on the presence of NP clusters in cells. The DP increases and the damage threshold decreases for target cells that may have clusters of nanoparticles compared with control samples having for which no clusters were found during EM examination. For example, the damage probability, DP, for specifically targeted cells, reached its absolute maximum of 100% at a fluence of 5 J/cm2, while for nonspecifically targeted cells it was only 0.07 at the same fluence. For the control sample of cells incubated with bare nanoparticles, the DP level was only 0.09 at a much higher fluence of 35 J/cm2. The damage threshold for the specifically targeted K562 cells is estimated at about 1-2 J/cm2, which is 30-50 times lower than the threshold fluence level for destruction of the same cells without particles (control). Thus, the specific targeting protocol provided a significant decrease of the laser damage threshold compared to intact untargeted cells and more than 10 times decrease compared with control K562 cells incubated with NPs not conjugated to a MAB. The length of PT response signal in targeted cells statistically varied from 0.2 to 3.5 ms. Maximal micro-bubble diameter reached 20 mm, as calculated using previously developed model of cell damage based on the experimentally measured lifetime of a microbubble [22].

Laser Elimination of Tumor Cells-Cell Suspension Model

In addition to laser irradiation of individual cells and their real-time monitoring with the PT microscope, suspensions of the human leukemic and normal stem cells were irradiated in the cuvettes, simulating the first approximation to the bone marrow purging procedure. In this case many cells (200-400) were irradiated simultaneously with a broad laser beam. Damaging effects of a single laser pulse was monitored by microscopic examination of trypan blue uptake by cells after the laser treatment. Cells with visual signs of destruction as well as positively stained were considered damaged. Laser fluence in the range of 0.5-2 J/cm2 was used in this experiment. At the level of 1.7 J/cm2, 100% of specifically targeted tumor cells were damaged (FIG. 10B) and 16% of normal stem cells were damaged (FIG. 9), while 84% of normal cells survived the laser pulses (FIG. 10A). The bubble generation threshold for tumor cells was in the range 0.1-0.3 J/cm2. This is 100-300 times lower than the bubble generation threshold of optical fluence (30-70 J/cm2) determined for intact untreated cells such as lymphocytes, K562 and lymphoblasts.

Example 2

Ex Vivo LANTCET of Human Normal and ALL Leukemia Cells Using Spherical Nanoparticles

Cryopreserved samples of human bone marrow (BM) taken either from normal donors or patients with the diagnosis of acute B-lymphoblast leukemia (ALL) were used. Suspensions of normal and tumor cells were not mixed. Normal BM samples had no tumor cells and ALL tumor samples consisted mainly of tumor cells where the level of B-lymphoblasts was 94% to 98% in samples from different patients. Three normal and three tumor samples were all obtained from different patients. Normal and tumor samples were prepared and analyzed as separate samples though the same treatment protocol was used for each. Leukemia cells express diagnosis-specific genes that were determined individually for each ALL patient by using standard clinical protocols [53]. Specific MAB, raised against cell membrane receptors corresponding to specific genes, were used for each sample of tumor cells. Phenotyping was performed with flow cytometry.

As a result diagnosis-specific (primary) monoclonal antibodies were determined (MAB1). For different patients different MAB yielded maximal levels of expression. It was determined that for each of the three patients an optimal MAB1 was different, i.e., CD10 for patient no. 1, CD19 for the patient no. 2 and CD20 for patient no. 3. These MABs were used and each patient-specific sample was treated with an MAB1 that corresponds to that sample.

LANTCET Method

High selectivity of formation of clusters of nanoparticles in the target tumor cells is achieved through a two stage incubation procedure. At the first stage the samples (normal and tumor) were separately incubated for 30 minutes at 4° C. with diagnosis-specific MAB1 (mouse anti-human CD10, CD19 or CD20 for each tumor sample) that selectively attach mainly to blast cells. Then both samples were separately incubated for 30 minutes at 4° C. with 30 nm gold spherical NP that were conjugated with secondary MAB2 (#15754, Ted Pella, Inc., Redding, Calif.). NP-MAB2 comes as a factory-made complex and MAB2 (goat anti-mouse IgG(H+L)(AH)) has high coupling efficiency for MAB1. The gold nanoparticles were attached to cell membranes through the chain of bonds, i.e., cell receptor-MAB1-MAD2-NP. All operations during stage 1 were performed at 4° C. which minimizes any physiological processes and allows efficient MAB-receptor or MAB-MAB interactions. Selectivity of the targeting after stage 1 is good, however, not sufficient from the clinical point of view. First, MAB1-specific receptors also may be expressed at minor levels in some normal cells. Second, MAB2 may directly couple through different uncontrollable mechanisms directly to the membranes of normal, i.e., non-target cells. Thus, small amounts of NPs may attach to membranes of normal cells causing death during exposure to laser radiation.

Therefore, selectivity and LANTCET safety is improved via a second incubation stage to create maximal clusters of NP inside the target cells. A temperature of 37° C. and incubation time of 30-120 minutes stimulates internalization of NPs through endocytosis into the cell endosomes. At the end of stage 2 large NP clusters were formed only in those cells, i.e., the target cells, with a high initial level of membrane-bound NP. Because of a much lower level of NP at the membranes of normal cells, the clusters either were not formed at all or had much smaller dimensions. The bigger the cluster size, the lower the laser fluence threshold of bubble generation. An explanation for such reduction of the laser fluence threshold is found in the combination parameters responsible for vapor bubble formation, i.e., larger NP volume increases heat accumulated in the vapor, longer heat diffusions time increases effectiveness of laser interaction with light-absorbing clusters, and a reduced surface tension and dynamic viscosity allows expansion of vapor nuclei into microbubbles [54].

Next, normal and tumor cell suspensions were injected into separate sample chambers with a diameter of 2.5 mm (S-24737, Molecular Probes, OR), and exposed to single laser pulses at a 532 nm wavelength matching a peak of light absorbance for spherical 30 nm gold NP. The laser pulse duration was 10 ns, which approximately matches the heat diffusion time from a 200 nm cluster of gold NPs into water. The laser energy(fluence) was set at the minimal level that provides bubble generation only around the biggest NP clusters (>200 nm) and does not induce any bubbles around smaller clusters or around single NPs which have a higher threshold for bubble generation. This fluence level provides high selectivity of cell killing because no bubbles can be generated in normal cells even if they accumulated some NP. Two experimental modes, i.e., single cell irradiation and simultaneous irradiation of many cells in suspension, were studied. Tumor and normal samples were exposed to laser pulses with equal fluence. In a single-cell mode (FIG. 11A) each individual cell (total of 150) was irradiated one by one with single focused laser pulse at 532 nm (LS 2132, Lotis TII, Minsk, Belarus). This mode was used for investigation of laser-induced bubbles in cells. Biological damage to irradiated cells was studied in the suspension mode (FIG. 11B) through simultaneous irradiation of 4,000-20,000 cells in a sample chamber with a wide laser beam (3 mm diameter, 532 nm, 10 nanoseconds). After exposure to laser pulses the samples were collected from the chambers and analyzed for cell viability.

Imaging and Measuring Nanoparticles in Individual Cells

CCD camera (model U2C-14S415, Ormins Ltd, Minsk, Belarus) with a 12-bit dynamic range and a sensor size of 1300×1000 pixels was used with a fluorescent light microscope (Leica DML, Leica Microsystems, Wetzlar, Germany) with 100×-microobjective. Standard “green” fluorescent excitation mode was applied. Fluorescent signal amplitude was acquired and measured in counts (0-12000) of camera digitizer. R-phycoerythirin (PE) fluorescent dye (#P9787), Ted Pella, Inc., Redding, Calif.) factory-conjugated with MAB3 (anti-goat IgG), was used as a marker for NP. PE was coupled to NP through MAB3-MAB2 bonds and by incubating the tumor and normal samples with PE-MAB3 after stage 1 and prior to stage 2 treatment.

Optical concentration of PE was determined by flow cytometry. Tumor cells were divided into two samples. Sample 1, NP-free cells, was incubated only with PE-MAB3 for 30 min at 4° C. and was used as a reference source of non-specific fluorescence of PE. Sample 2 was treated according to stage 1 protocol (cells with NP) and then was incubated with PE-MAB3 for 30 min at 4° C. Sample 2 was used as the source of Np-specific fluorescence. Several concentrations of PE were used. Concentration of PE 1:2,5000 provided maximal ratio of 18 for the mean amplitudes of flow cytometer output signals of the sample 2 (51.9±29.5 au) to the sample 1 (2.85±2.58 au). This PE concentration was used in all experiments for detecting and imaging NP in cells and the level of non-specific fluorescence was less than 6%.

To allow quantitative estimation of NP concentration in individual cells, the CCD camera was calibrated. The images of homogenous NP solutions (without cells) in water at several different concentrations of NP were obtained. The mean pixel amplitude S of fluorescent signal was found to be almost linearly proportional to the concentration of NP, which allowed one to estimate the NP local concentration C through fluorescent amplitude S in the corresponding point of the image by:


C=0.022×1011×(S−Sbc−Sns)  (11)

where Sbc is a background signal of image detector (459 counts), Sns is a level of non-specific signal (6%).

For each cell, optical and fluorescent images were obtained prior to its exposure to a laser pulse. Optical image was used for determining cell diameter and location of outer membrane and nuclear membrane. The fluorescent image was analyzed with special software that was previously developed as part of the photothermal microscope (21). Several image parameters were used to analyze such as NP-related properties as the mean level of NP in an individual cell, which is measured as the cell-averaged pixel amplitude, the amount of NP in clusters which characterizes cluster size and is measured as the maximal amplitude in the center of cluster-related image, i.e., Max, the spatial radial distribution of NP and their clusters inside the cell which are measured as the mean relative radius Mir (normalized by cell radius) corresponding to location of maximal amplitude, and the heterogeneity of spatial distribution of the Np. Image parameters were calculated for each cell and then were analyzed as histograms for each sample.

For the images with size close to the diffraction limit of the microscope (300-350 nm), an actual size of the source of fluorescent signal can not be determined because it can be much less than that. Thus, the peak image amplitude was treated as the measure of total number of NP in one cluster. Conventional (non-confocal) microscopic image is the two-dimensional projection of three-dimensional distribution of fluorescent signals in the depth of focus of the micro-objective. Still, membrane-type fluorescence could be differentiated from cytoplasm-type fluorescence and the uniform fluorescence from cluster-type fluorescence.

Detection of Laser-Induced Microbubbles in Cells

The photothermal microscope was used for detecting vapor bubbles in individual cells. Each cell was illuminated with three collinear focused laser beams, i.e., a pulsed beam for bubble generation and two low-intensity continuous (633 nm, 0.1 mW) and pulsed (640 nm, 8 ns) probe beams for detecting the bubble in a cell in real time. Any bubble-related change of refractive index causes the shift of the phase of the probe beam that influences beam intensity in the input of the photodetector. Output of the photodetector is measured as PT response by a high speed digitizer (Bordo-211, Auria Ltd., Belarus) in case the bubble PT response has a specific shape so that bubble generation can be detected during irradiation of individual cells with a laser pulse.

Each individual cell (total of 150) was irradiated one by one with a single focused laser pulse at 532 nm (Lotis TII, Minsk, Belarus) and the PT response from each cell was simultaneously recorded as the time-response of the probe laser intensity [22]. This mode allowed both bubbles detection and lifetime measurement that characterizes maximal bubble diameter. After sequentially irradiating 150 cells with a single pulse and registering the PT response from each one, the probability PRB of bubble generation at a specific laser fluence was as described in Example 1.

For time-resolved imaging of the bubbles in individual cells, a dye pulsed probe laser beam at the wavelength of minimal light absorption by the cell (640 nm, 8 ns) was used and which was synchronized and delayed for 120 ns against the pump pulse [55]. Bubble location and diameter were analyzed from probe beam images captured with a CCD camera so that only events occurring after this delay and during probe pulse duration were visualized.

Cell Damage Detection

Laser-induced damage was measured by microscopy (concentration counts in hemocytometer) and flow cytometry (outer membrane damage detection as propidium iodide, PI, positive stains) after each stage of sample preparation and 2 hours after laser treatment. Microscopy detected the destruction of cells and flow cytometry detected necrotic death of those cells that were damaged, but not destroyed. Cells with compromised membranes that included propidium iodide (PI) were considered as damaged. The cells that were counted as PI-negative were considered as survived live cells. The level of survived live cells LLC was used as a measure of LANTCET efficacy and was analyzed experimentally for each normal and tumor sample as a function of laser pulse fluence, number of laser pulses, types of MAB1, and incubation stage.

This parameter describes the relative change of the level of living cells in population due to laser treatment and counts cell losses due to destruction and cell death:


LLC=(Cal/Cbl)*CPI−*100%  (eq. 12)

where Cbl is the initial concentration of all cells in the sample (counted in hemocytometer) before laser treatment, Cal is the concentration of all cells in the sample after laser treatment (both counted in hemocytometer), CPI− is the level of PI-negative (live) cells obtained with a flow cytometer for MAB1-positive cells in the tumor sample and for all cells in the normal sample. Delayed cell death was not considered because the bubble-related mechanism of cell damage acts very fast.

Fluorescent Imaging of Nanoparticles in Cells

Fluorescent images of tumor cells showed the two types of spatial distribution of the signals: local peaks and uniform areas (FIGS. 12A-12F). Local and relatively strong fluorescent peaks were found in most of the images of the tumor cells that were pre-incubated with nanoparticles. These peaks were associated either with cell membrane (FIGS. 12C, 12E) or with cell cytoplasm (FIGS. 12D, 12F). The amplitude in the peak (2000-10000 counts; FIGS. 12E-12F) was 5-20 times higher than that for cell areas with uniform signal (200-500 counts). All peak-related images had a similar shape, a round spot with the diameter 0.4-0.8 μm, that is close to the diffraction limit of the microscope. The fluorescent signals of such shape and amplitude were never observed in cell-free space. Such peaks are considered as evidence for the clusters of nanoparticles.

Using calibration data, the concentration of single NP in tumor cells was estimated as 11.2×1011±4.7×1011 for the uniform areas, while in the clusters it varied from 57×1011±21×1011 (48 C, 0.5 hours) to 122×1011±55×1011 (37° C., 2 hours). Concentration of NP in tumor cells was found to be approximately 10 times (uniform areas) and 100 times (clusters) higher than the concentration of free unbound NP during the first stage of cell incubation. The actual size of NP clusters could not be measured because it may be below the optical diffraction limit of the microscope and, therefore, the actual concentration of NP in the clusters may be even higher. Nevertheless the maximal peak amplitude characterizes the total number of NP in the cluster. Some of the normal BM cells also exhibited fluorescent images (not shown) with cluster-related peaks that were similar to those obtained for the tumor cells (FIG. 12C). Although the total level of NP-positive normal BM cells was within 6%, the rest of normal cells did not yield NP-specific fluorescence in their images.

Nanoparticle Accumulation in Cells after Stage 1 and Stage 2

The peak amplitudes and the homogeneous fluorescence were measured in tumor and normal cells as a function of the incubation time after stage 1 (4° C., membrane coupling) and stage 2 (37° C., internalization). All fluorescent images were analyzed in two ways. First, the image parameters were calculated for each cell and were plotted as the histograms for each stage of cell incubation (FIG. 13A). Second, cell counts were performed for three categories of fluorescent images, i.e., images without NP clusters, images with peripheral NP clusters only that are located at cell outer membrane and images with peripheral and intracellular (located inside cell) clusters of NP (FIG. 13B). Image parameters and cell counts were used to analyze the influence of incubation temperature and time on clusterization of NP.

Distinct intracellular clusters of NP were found only for the tumor cells that were incubated at 37° C. (after stage 2, see FIG. 11D, 11F) with fluorescent peaks having maximal amplitudes of 8,000-11,000 counts. To compare NP cluster size after stage 2 and stage 1 the histograms were analyzed for the maximal values of fluorescent amplitudes in peaks (FIG. 13A). This image parameter is important because the largest clusters of NP in the cell would produce the biggest bubbles and thus such clusters are the main cell-killing nanostructures. Maximal peak amplitudes after stage 1 were 2,000-4,000 counts for all incubation times, which is two to four times lower than those after stage 2 (2 hours of incubation). This difference may be interpreted as the difference in NP cluster size because the amplitude is proportional to the total number of NP in one cluster and the size of the cluster correlates to the number of NP in it. Therefore the stage 2 (incubation at 37° C., 2 hours) provided the largest clusters of NP with their localization inside cells. Fluorescent peaks within cell nuclei were not observed. Thus, the internalization of NP was confined by the cytoplasm; NP did not penetrate into cell nuclei that are quite large for this type of cells (FIG. 11A-11D). Spatial distribution of fluorescent peaks in cells was quantified through corresponding image parameter Mir and as function of the time and temperature of cell incubation (FIG. 13B). Histogram analysis of this image parameter showed a clear trend in re-localization of NP clusters from the periphery (Mir value 0.6-0.8) after stage 1 to the inner site of the cell (Mir value 0.3-0.6) after stage 2 of incubation. Such spatial distribution of the fluorescent peaks confirms endocytotic mechanism of NP clusterization.

Additional measurements were made to understand the kinetics of the internalization process. The cells with images were counted as in the three categories defined above categories, i.e., the number of cells with intracellular peaks in their images and the number of cells with membrane-located peaks in their images, as function of cell incubation time and temperature (FIGS. 14A-14B). No changes were found in those categories for normal BM cells (FIG. 14A) regardless of the time and the temperature of cell incubation. The counts of cluster-related cells were within 6.0%±1.1% for any incubation conditions. So stage 2 did not add any NP or NP clusters to normal cell after stage 1. For tumor cells the situation was totally different (FIG. 14B). Incubation at 37° C. (stage 2) caused steady increase of the number of cells with intracellular clusters from 12.0%±2.0% at the beginning of stage 2 to 66.0%±4.0% after 2 hours. Also during stage 2 tumor cells yielded the decrease of cell counts for membrane-located clusters from 84.0%±4.2% at the beginning to 30.0%±2.8% after 2 hours. Tumor cells that were incubated at 4° C. (stage 1) showed no changes at all in the counts of membrane-located and intracellular clusters (FIG. 14B). The level of intracellular NP clusters after stage 1 was under 8.0%±0.9% and did not increase for 2 hours. Obtained experimental results demonstrated that the formation of clusters of NP occurred at the cellular membrane and inside the cells and that the largest NP clusters emerge inside in the cytoplasm of the tumor cell during the second stage of incubation from internalization. The difference in NP levels in normal and tumor cells indicated significant increase in selectivity of NP clusterization in tumor cells in comparison with normal cells after an additional second stage of incubation.

Laser-Induced Bubbles in Cells

Photothermal (PT) responses and images were obtained for individual tumor cells (single cell mode) at different laser fluencies. Typical bubble-specific PT responses and images that were obtained after a single laser pulse at 0.6 J/cm2, 532 nm for incubation conditions of 37° C., 2 hrs are shown in FIGS. 15A-15C. The duration of the PT response allows measurement of the bubble lifetime and thus its maximal diameter can be estimated. Location of the bubble-specific signals in the PT image of the tumor cell (FIG. 15A) shows that the bubbles were generated in the cytoplasm and probably not in the area of the nucleus. The bubble generation sites spatially coincide with location of nanoparticle clusters as shown in the fluorescent images. Also it was found that the number of bubble-related signals in one cell, from 1 to 4, was lower than the number of cluster-related peaks, 4-20, in those cells. Therefore, under a given laser fluence of 0.6 J/cm2, not every cluster produced the bubbles. It is likely that only biggest clusters of nanoparticles are cell damaging agents while the rest of smaller clusters and single nanoparticle may not produce the bubbles and therefore do not contribute to nanothermolysis of the cell.

The PT response of the bubble (FIG. 15A) was obtained at the fluence of 0.6 J/cm2. No bubble-specific response was detected at the same fluence when water suspension of individual nanoparticles at a concentration of 8×1011/ml was irradiated with a single laser pulse. Also, such PT responses were not detected for normal bone marrow cells. This result may be considered as additional evidence that only clusters of nanoparticles produced the bubbles at this fluence level. It also was discovered that the cells with nanoparticle clusters, shown in their fluorescent images, may generate the bubbles during exposure of the cell to several (up to 60) laser pulses. Single NPs that produced bubbles at much higher fluencies exhibited first-pulse bubble generation though never produced the bubbles during the second and following pulses because they were destroyed, i.e., melted or evaporated, during the first laser pulse. Thus, the cluster of nanoparticles are a photostable structure, unlike the single nanoparticles that are destroyed after the first pulse, and can be irradiated with more than one laser pulse. Using the model of a laser-activated bubble, maximal bubble diameter is estimated as 13 μm for the PT signal (FIG. 15C). Such a bubble size is comparable with the cell diameter, that is, 8-9 μm (FIGS. 11A-11B) and therefore may rupture the cell outer membrane. This causes lysis of any tumor an normal cell.

Laser-Induced Damage to Cells

Cell samples were irradiated in round 2.5 mm cuvettes with single or several laser pulses at 532 nm. Cell damage was analyzed through LLC dependence upon laser parameters (pulse fluence and number of pulses) and incubation parameters (nanoparticle concentration, incubation temperature and MAB1 types). Influence of the laser parameters on the cell damage was studied for the same incubation conditions such as one type of MAB1 (specific for each patient), incubation temperature of 4° C. and a nanoparticle concentration during the first stage of incubation of 15000 nanoparticle/cell.

Increase of laser fluence from 0.2 to 2 J/cm2 caused a gradual decrease of LLC for tumor cells from 3.9%±0.6% at 0.2 J/cm2 to less than 1% at the fluencies above 0.6 J/cm2. The samples obtained from 3 different patients with an ALL diagnosis showed different degrees of damage of tumor cells. At the fluence of 0.6 J/cm2 and single-pulse irradiation, the LLC was found to be 1.5%±0.3%, less than 0.1%, and 5%±0.6% respectively for the patients 1, 2 and 3. These samples were treated with different MAB1 and the difference detected in LLC values may be caused by the variations in nanoparticle targeting efficacy. The change in LLC of the tumor cells after laser treatment was due mainly to the decrease in the concentration of cells, i.e., up to 10 times in comparison with initial concentration, and, to a lesser extent, to cell membrane damage. The LLC for normal bone marrow cells under the same conditions was found to be 77%±6.2% to 84%±4.2%. Irradiation of tumor cells with 10 laser pulses instead of 1 did not produce a significant effect. The LLC decreased from 1.5%±0.3% to 1.2%±0.2%. This means that the first laser-induced bubbles produce the damage and unlike the damage through heating the bubble-related damage does not have accumulative nature and occurs immediately after expansion of the first bubble.

The influence of incubation parameters on cell damage was studied. Under a fixed concentration of nanoparticles during the first stage of incubation, the application of the different primary MABs resulted in a variation of LLC from 0 to 67% (FIG. 16A). The strongest effect, that is when LLC is 0, was reached by using the combination of several primary monoclonal antibodies. The influence of another incubation parameter, the temperature of the second stage of incubation, is shown in FIG. 16B. Increase of the incubation temperature from 4° C. to 37° C. caused a 2.6× decrease of LLC for tumor cells from 3.9%±0.6% at 4° C. to 1.5%±0.3% at 37° C. and a 1.6× decrease of LLC for normal cells from 77%±6.2% at 4° C. to 49%±5% at 37° C. Therefore in terms of the efficacy of cell killing the two-stage incubation at 37° C. has produced better results than the standard incubation scheme at 4° C. Although “hot” incubation was less safe for normal BM cells. No significant variation of LLC for tumor cells was discovered when the concentration of nanoparticles varied from 6000 to 150000 nanoparticles/cell during the first stage of incubation. The LLC of tumor cells decreased from 15% (6000 nanoparticle/cell) to 10% (150,000 nanoparticle/cell). It is assumed that the concentration from 15000 to 30000 nanoparticles/cell is quite saturating, because further increase of nanoparticle concentration had no effect on cell killing efficacy.

Example 3 LANTCET of Human Normal and Tumor Cells Using Nanorods or Nanoshells

Cells and Antibody Conjugates

K562 cells that have high level of expression of CD33 antigens were used. Primary samples of human bone marrow (BM) taken either from normal donors or patients with the diagnosis of acute myeloid leukemia (AML) were used in experiments. All AML (tumor) samples consisted mainly of tumor cells where the level of B-lymphoblasts was 94 to 98% in samples from different patients. AML cells expressing diagnosis-specific CD-33 genes were determined with flow cytometry.

Conjugates of gold nanorods (NR) with and without antibodies to CD33 were used. AML and K562 cells were incubated for 30 min at 37° C. with the gold nanorods with dimensions 45×14 nm and absorption maximum at 780 nm. Then living and fixed cells were used as suspensions for bubble generation and detection.

Solid tumor cells were prepared as the monolayers of living EGF-positive carcinoma cells (Hep-2C) that were grown on glass surface and human lymphocytes were used as normal cells. Both Hep-2C and normal lymphocytes were incubated 35 min at 37° C. with the conjugates of 40 nm gold nanoshells-C225 antibody that was grown against EGFR. Cells were irradiated with single laser pulse at the wavelength of 720 nm close to the peak of maximal optical absorbance of the nanoshells.

Generation and Detection of Photothermal Bubbles (PTB) and Imaging

The photothermal microscope described herein is used for generating and detecting PTB in individual cells. Each cell was illuminated with 3 collinear focused laser beams: pulsed beam for bubble generation (532 nm or 720-840 nm), and low-intensity continuous (633 nm, 0.1 mW) and pulsed probe beams for detecting the bubble in a cell in real time. Each individual cell (total of 90 for each sample) was irradiated one by one with single focused laser pulse of the same fluence at 532 nm (gold nanospheres) or 720-840 nm (gold nanoshells and nanorods) and PT image and response from each cell were simultaneously recorded as time-resolved image of pulsed probe laser and time-response of c.w. probe laser intensity. After sequential irradiating of all cells in the sample and registering PT images and responses from each cell we have measured the probability PRB of bubble generation at specific laser fluence is measured as described herein. A CCD-camera as described herein is used for time-resolved imaging of the bubbles in individual cells irradiated with a pulsed pump laser beam.

Optical Scattering Images and Spectra of Nanorod Clusters in Cells

Formation of nanorod (NR) clusters was verified directly with optical scattering microscopy and microspectroscopy. As the source of light we used A white light continuous source was used. Light was directed with an optical fiber at the glass slide with the sample at an angle of 70-80°. NR-treated and control (untreated) K562 cells were imaged with light scatter photothermal microscopy. Images of individual cells are shown in FIGS. 17A-17C. Those images revealed two common features both for the AML and K562 cells. NP-related signals were detected in 90-95% of targeted cells and spatial distribution of NP-related signals within the cell was highly heterogeneous with apparent local peaks. Also no NP-related signals were found in the images of control cells. Thus, regardless of the targeting method used the nanorods aggregate into clusters during interaction with living cells.

Amplitudes of scattering signals are 1230±580 for intact control cells and 3980±1170 for NR-treated K562 and AML cells. Scattering images were registered with a continuous white light source and no photobleaching effects occurred, so they are more suitable for NP imaging and quantitative studies. Untreated (control) cells also produced some optical scattering though with much lower amplitudes and without strong local peaks (FIG. 17A). Optical scattering sensitivity may be improved by using a monochromatic light source at the wavelength that matches plasmon resonance wavelength for NR clusters. Spectral studies of the cells were performed with a microspectrometer. Scattering spectra of a local zone with diameter not bigger than 1 μm were measured for control and treated cells (FIG. 17C). For treated cells the spectra were measured for the regions occupied by the clusters. The cluster-related spectrum demonstrated a resonant nature of scattering and is very close to that obtained for the water suspension of nanorods. Their maximums almost coincide (680 nm for NR water suspension and 670 nm for NR clusters in cells). The conclusion is that the local scattering signal represent the clusters of gold nanorods.

Photothermal Bubbles Around Single NP and in AML Cells Around Gold Nanorods

A typical bubble-specific PT response (PTB) and images are shown in FIGS. 18A-18C that were obtained for individual AML cells with nanorod clusters at different laser fluencies. Peripheral location of the bubble-specific signals in the PT image of the tumor cell shows that the bubbles were generated in the cytoplasm and probably not in the area of the nucleus. For all types of studied cells, including cultured and primary tumor cells, a single laser pulse accompanied by PTB caused cell lysis through mechanical damage of the cytoplasmic membrane. Cell membrane damage can be seen in its optical image (FIG. 18D). Even the smallest bubbles with lifetimes of 20-40 ns caused cell damage. Table 2 gives bubble generation thresholds for AML cells (780 nm, 10 ns).

TABLE 2 Bubble generation Samples thresholds, J/cm2 Intact cells >50 Cells with NR (4° C.) ≈1.5 Cells with NR (37° C.) 0.32

Analysis of the thresholds of PTB as shown in the Table 2 indicates that PRB increases gradually up to 100% with increase of pulse fluence and this means that ALL the cells in population accumulated nanorods. Also this dependence allows the bubble generation threshold to be determined as one value for all cells. Increasing the temperature during the incubation of the cells with nanorods from 4° C. (only chemical reactions are allowed) to 37° C. (all physiological reactions including endocytosis are allowed) caused significant increase of PRB and bubble size and the decrease of bubble generation threshold by 5 times. Also in the both cases PTB threshold levels are much lower than the threshold obtained for suspension of single nanorods in water. The conclusion in both cases is that clusters of nanorods were formed and that nanorod clusters formed at 37° C. are larger.

Spectral dependence of PTB lifetime and probability of PTB generation around single NR in their water suspension and after their clusterization in human bone marrow CD33+ AML cells was measured (FIGS. 19A-19B). Bubble generation probability (PRB) was measured at several NIR wavelengths at fixed laser fluence levels that was 7.5 J/cm2 for single nanorods in water and 0.75 J/cm2 for the cells treated with the same nanorods. The PRB spectrum for single nanorods closely matches its absorption spectrum, while formation of nanorod clusters has broadened the spectrum of PRB obtained for the cells.

The PRB spectra for nanorods in water matches optical absorption spectra of the same nanorod suspension with peak PRB and OD at 780 nm. This means that bubbles are generated due to absorption of laser energy. The cell spectrum for PRB also has peak around 750-780 nm and matches the spectrum of PRB for NR suspension. Significant broadening of PRB spectrum was not found as it was found for optical absorption spectrum of aggregated NR in water. Notably, the absorbance and PRB spectra of the nanorod suspension in water match each other in terms of width of peak, but in the case of cells there is no such match. Absorbance spectrum of suspension of nanorod clusters is very broad, but PRB spectrum of cells that apparently have nanorod clusters is not so broad and the width of its peak is closer to the width of the peak of the PRB spectrum obtained for individual nanorods in water. Comparing bubble generation thresholds obtained for single spherical nanoparticles and nanoshells in water, for normal cells and tumor cells indicates that target tumor cells form big clusters resulting in the decrease of bubble generation/cell damage laser fluence threshold in tumor cells by almost 100 times relative to untreated cells. Table 3 gives laser fluence thresholds (J/cm2) for generation of PTB in normal and ALL tumor cells.

TABLE 3 Gold red Single NP NP type lymphocytes blood cells K562 AML In water Spherical NP 9.4 27.0 43 (bare, 30 nm) (532 nm) (532 nm) (532 nm) NP-IgG-CD33 0.3 0.42 (conjugates) (532 nm) (532 nm) NR-PEG  0.43 0.32 17 45 × 14 nm (780 nm)  780 nm)  780 nm) No particles 10.0 2.6 >45    >35    — Intact cells (532 nm) (532 nm) (532 nm) (532 nm) />40   (780 nm) />50   />50    />50   (780 nm) (780 nm) (780 nm)

Photothermal Bubbles Formed in Solid Tumor Cells Around Gold Nanoshells

A solid tumor application was studied with Hep-2C model cells and conjugates of 40 nm gold nanoshells with EGF-specific antibody C225. Bubble images and responses obtained from carcinoma cells Hep-2C are shown in FIGS. 20A-20D. Comparing bubble generation thresholds obtained for single nanoshells in water, for normal cells, and for tumor (Hep-2C) cells showed that normal cells uptake some single nanoshells that do not form clusters and that target tumor cells form large nanoshell clusters. These large clusters cause a decrease in the bubble generation/cell damage laser fluence threshold in Hep-2C cells by almost 100 times, relative to untreated cells, and by 20 times, relative to normal cells that also were treated with NP-C225 conjugates and laser pulsed under the same conditions. Table 4 gives the laser fluence thresholds (J/cm2) for generation of PTB (laser pulse 720 nm, 10 ns) in normal and solid Hep-2C tumor cells.

TABLE 4 red Single NP NP type lymphocytes blood cells Hep-2C in water NS (bare) — 11.0 — NS-C225 12.0 — 0.5 11.0 (conjugate) No particles >40 >50 >40 — Intact cells

Example 4 LANTCET of a Sarcoma in a Rat Model Using Nanoparticles

Rats were used to grow polymorphic sarcoma 1 (tumor type M-1 obtained from the bank of Russian Oncological Research Center) to a diameter of 0-15 mm. The skin layer was removed at the tumor site. A drop of a nanoparticle suspension (5 μl) was administered onto the tumor surface and left for 40 min to be absorbed by the tumor. Gold spherical 30 nm particles conjugated to goat anti-mouse IgG (#15754, Ted Pella, Inc, Redding, Calif.) were then applied topically onto the tumor surface. The skin layer was removed at the tumor site. The drop of spherical nanoparticles suspension (5 μl, with concentration of NP 8×1011/ml) was administered onto the tumor surface and left for 40 min to be absorbed by the tumor at room temperature. The tumor site was then treated with the laser. A single laser pulse at 532 nm (maximum of light absorbance by NPs), 10 ns duration, fluency of 0.75 J/cm2, and diameter of 3-4 mm was directed to the central area of the tumor. After 24 hours following the laser treatment the solid tumor was extracted from the animal and the degree of tumor necrosis was measured according to the uptake of trypan blue.

The tumor that was treated with a single laser pulse showed a necrotic (FIG. 21A, white area) area with diameter close to the laser beam diameter (3-4 mm) and the depth of 1-2 mm that indicates the spherical nanoparticle diffusion depth. The tumor that underwent laser treatment without pretreatment with NPs showed no signs of necrosis (FIG. 21B). The conditions of spherical NP-cell interaction at physiological temperature and 40 min of interaction time were sufficient to allow NP cluster formation inside tumor cells.

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Any publications or patents mentioned in this specification are indicative of the levels of those skilled in the art to which the invention pertains. Further, these publications are incorporated by reference herein to the same extent as if each individual publication was specifically and individually incorporated by reference.

One skilled in the art will appreciate readily that the present invention is well adapted to carry out the objects and obtain the ends and advantages mentioned, as well as those objects, ends and advantages inherent herein. Changes therein and other uses which are encompassed within the spirit of the invention as defined by the scope of the claims will occur to those skilled in the art.

Claims

1. A method for increasing selective therapeutic thermomechanically induced damage to a biological body, comprising:

specifically targeting a biological body comprising a medium with a plurality of nanoparticulates each conjugated to at least one targeting moiety, said nanoparticulates effective to form one or more nanoparticulate clusters on or in the biological body upon targeting thereto;
irradiating the biological body with at least one pulse of electromagnetic radiation having a spectrum of wavelengths selected to have a peak wavelength that is near to or that matches a peak absorption wavelength of the nanoparticulates; and
generating vapor microbubbles from heat produced via absorption of the electromagnetic radiation into the nanoparticulate cluster(s), wherein the vapor microbubbles cause selective and increased thermomechanical damage to the targeted biological body.

2. The method of claim 1, further comprising:

filtering the products of the thermomechanical damage from the medium.

3. The method of claim 1, further comprising:

receiving a photothermal signal or generating an optical image of thermomechanical effects to monitor and to guide selective thermomechanical damage to the biological body.

4. The method of claim 1, wherein the nanoparticulate has a dimension of about 1 nm to about 1000 nm.

5. The method of claim 1, wherein the nanoparticulate cluster has a total volume about 2 to about 200 times greater than a volume of the nanoparticulate comprising the same.

6. The method of claim 1, wherein the nanoparticulate is a spherical nanoparticle, a nanorod or a nanoshell at least partially comprising gold or silver or is a carbon nanotube.

7. The method of claim 1, wherein the targeting moiety is a monoclonal antibody or a peptide specifically targeted to a receptor site on the biological body.

8. The method of claim 7, wherein the receptor site further comprises another monoclonal antibody or peptide attached thereto specific for the targeted monoclonal antibody.

9. The method of claim 7, wherein the nanoparticulate further comprises complementary strands of a nucleic acid conjugated thereto or a combination thereof.

10. The method of claim 7, wherein the nanoparticulate further comprises PEG molecules.

11. The method of claim 1, wherein the wavelength spectrum is a range of wavelengths of about 300 nm to about 300 mm.

12. The method of claim 1, wherein the pulse of electromagnetic radiation is optical radiation.

13. The method of claim 12, wherein the pulse of optical radiation has wavelength in the range from 500 nm to 1150 nm.

14. The methods of claim 1, wherein the pulse of electromagnetic radiation is about 1 ns to about 100 ns in duration.

15. The method of claim 1, wherein the biological body is an abnormal cell, a bacterium or a virus.

16. A system for increasing selective therapeutic thermomechanical damage to abnormal cells, comprising:

a chamber containing the abnormal cells in a medium;
a source of nanoparticulates modified to specifically target the abnormal cells fluidly connected to the cell chamber;
an optical chamber adapted to contain the targeted abnormal cells fluidly connected to the cell chamber;
a pulsed source of electromagnetic radiation directed against the targeted cancer cells in the optical chamber, said source configured to emit a spectrum of wavelengths selected to have a peak wavelength that is near to or that matches a peak absorption wavelength of said nanoparticulates; and
means for filtering out cells damaged by thermomechanical effects resulting from absorption of the electromagnetic radiation emitted at the peak wavelength, said filtering means fluidly connected to the cell chamber.

17. The system of claim 16, further comprising:

means for receiving a photothermal signal or for generating an optical image of the thermomechanical effects.

18. The system of claim 16, wherein the nanoparticulates each comprise at least one targeting moiety specifically targeted to a receptor site on the cancer cell.

19. The system of claim 18, wherein the receptor site further comprises another targeting moiety attached thereto specific for said targeting moiety on the nanoparticulates.

20. The system of claim 18, wherein the targeting moiety is a monoclonal antibody or a peptide attached thereto specific for the targeted monoclonal antibody.

21. The system of claim 18, wherein the nanoparticulate further comprises complementary strands of a nucleic acid conjugated thereto or a combination thereof.

22. The method of claim 21, wherein the nanoparticulate further comprises PEG molecules.

23. The system of claim 16, wherein the nanoparticulate has a dimension of about 1 nm to about 1000 nm.

24. The system of claim 16, wherein the nanoparticulate is a spherical nanoparticle, a nanorod or a nanoshell at least partially comprising gold or silver or is a carbon nanotube.

25. The system of claim 16, wherein the wavelength spectrum is a range of wavelengths of about 300 nm to about 300 mm.

26. The system of claim 16, wherein the pulse of electromagnetic radiation is optical radiation having a wavelength in the range from 500 nm to 1150 nm.

27. The system of claim 16, wherein the pulse of electromagnetic radiation is about 1 ns to about 100 ns in duration.

28. The system of claim 16, wherein the abnormal cells are leukemic cancer cells.

29. A method for treating a leukemia in an individual, comprising:

a) obtaining a sample comprising normal and leukemic cells from the individual;
b) placing the sample in the cell chamber of the system of claim 16;
c) targeting the cancer cells in the sample with the modified nanoparticulates, said modified nanoparticulates forming one or more clusters on or in the targeted cancer cell;
d) irradiating the targeted leukemic cells with electromagnetic radiation emitted from the pulsed source, wherein the electromagnetic radiation absorbed by the nanoparticulates causes selective and increased thermomechanical effects damaging to the targeted cancer cells, but not to the normal cells comprising the sample;
e) filtering out the damaged cells from the sample;
f) returning the normal cells remaining in the sample to the individual thereby treating the leukemia; and.
g) repeating said method steps a) to f) zero or more times, thereby treating the leukemia.

30. The method of claim 29, further comprising:

receiving a photothermal signal or generating an optical image of the thermomechanical effects to monitor and to guide selective thermomechanical damage to the cancer cells.

31. The method of claim 29, wherein the thermomechanical effects are caused by heat generated within the nanoparticulates from absorbed electromagnetic radiation sufficient to form vapor microbubbles around the nanoparticulate clusters.

32. The method of claim 29, wherein the nanoparticulate cluster has a total volume about 2 to about 200 times greater than a volume of the nanoparticulate comprising the same.

33. A method for selectively and thermomechanically damaging cells associated with a pathophysiological condition, comprising:

targeting the cells with a first monoclonal antibody specific thereto;
targeting the cells with gold nanoparticulates modified with a second monoclonal antibody specific to the first monoclonal antibody whereupon one or more clusters of the nanoparticulates form on or in the targeted cells;
heating one or more clusters of gold nanoparticulates formed on or in the targeted cells; and
generating vapor bubbles around the heated clusters sufficient to thermomechanically damage the cells.

34. The method of claim 33, further comprising photothermally or optically monitoring the thermomechanical damage.

35. The method of claim 33, wherein the gold nanoparticulates are spherical nanoparticles, nanorods or nanoshells.

36. The method of claim 33, wherein the clustered gold nanoparticulates are heated with optical radiation having a wavelength in a range from 500 nm to 1150 nm.

37. The method of claim 33, wherein the optical radiation is pulsed for a duration of about 1 ns to about 100 ns.

38. The method of claim 33, wherein the nanoparticulate has a dimension of about 1 nm to about 1000 nm.

39. The method of claim 33, wherein the nanoparticulate cluster(s) has a total volume about 2 to about 200 times greater than a volume of the nanoparticulate comprising the same.

40. The method of claim 33, wherein the cell is a cancer cell, a bacterial cell or a virus.

Patent History
Publication number: 20120046593
Type: Application
Filed: Aug 15, 2011
Publication Date: Feb 23, 2012
Inventors: Alexander Oraevsky (Houston, TX), Dmitri Lapotko (Minsk)
Application Number: 13/136,939