Patent Application
20090304587
The present invention generally relates to composites for internal local radiotherapy.
Radiotherapy is the medical use of ionizing radiation for the treatment of a variety of disorders, among which cancer treatment is the most widespread. Radiotherapy is used for either curative or adjuvant treatments of cancer. However, it is often used palliativly to accomplish local control over metastatic spread and it is most common to blend radiotherapy with surgery, chemotherapy, hormone therapy and combination of thereof.
Because radiotherapy is applied to the gross tumor and marginal normal tissues (and sometimes to neighboring draining lymph nodes), healthy tissues are often damaged by the high energy of the external beam used (external beam therapy). One way to reduce the injury is to use shaped radiation beams, from different angles (at a distance of 50 cm to several meters) to intersect the tumour, a tactic that provides a larger radiation dose than in the surrounding, healthy tissues.
The alternative approach is the use of brachytherapy. In brachytherapy (“short distance therapy”) the radioactive source (metal seeds or ribbons) is implanted either within (interstitial implants) or in close proximity (intra-cavitary implant) to the tissues of interest or in contact with the tissue at risk. In accord with the tumor type brachytherapy is categorized as (a) mould brachytherapy for the treatment of superficial tumors (skin); (b) surface brachytherapy, where a tinny applicator (hollow thin silver casting, containing a radioactive source) is placed on the ill organ; (c) interstitial brachytherapy, where radioactive sources (metal needles) are inserted into the tissue (e.g. prostate); (d) intracavitary brachytherapy, in which the sources is implanted inside a pre-existing body cavity and (e) intravascular brachytherapy, where a loaded catheter is placed inside the vasculature (in-stent restenosis).
Permanent implants using radioactive “seeds” containing various radiation sources such as 125I have been employed. 137Cs, 192Ir, and 103Pd sources have been employed in temporary implants. The use of 133Xe and 131Xe has also been suggested.
European Patent No. 0979656 discloses a radioactive composition for use in the therapy of neoplastic and non-neoplastic diseases by means of application of radioactive sources in direct contact with the tumor tissues or within them, and for use as an aid in radio-guided surgery. The composition contains radioisotopes immobilized on biocompatible or bioabsorbable solid particles, incorporated in a biocompatible and bioabsorbable matrix consisting of a hypotonic gel containing 2 to 30% w/w polyvinylpyrrolidone and 0.01 to 2% w/w of agar-agar or agarose in water.
International Patent Publication No. WO 97/19706 discloses therapeutic sources for use in the practice of brachytherapy. They are made of radioactive powder of 103Pd, 192Ir, 90Yt, 32P or 198Ag dispersed in a biocompatible polymeric matrix. The polymeric matrix is desirably manufactured with pre-selected flexibility suitable to its intended use, e.g. in the form of a rod, hollow rod, suture, film, sheet, or microspheroidal particles. The radioactive composites generate a substantially uniform radiation field in all directions. The radiation dose is assembled by from the radioactive composite during the medical procedure to emit a desired amount of therapeutic radiation. Optionally, the polymer is selected to dissolve or degrade in the body at a predetermined rate, depending upon the half-life of the radioisotope used in the therapeutic source.
International Patent Publication No. WO 95/16463 discloses the use of therapeutic radioactive compositions comprising radioactive elements such as 166Ho, 153Sm, 105Rh, 177Lu, 192In, 165Dy, 90Y, 140La, 159Gd, 175Yb, 186Re, 188Re and 47Sc, adsorbed onto cellulose ether or a derivative thereof in treating rheumatoid arthritis.
While brachytherapy has proven safer for adjacent healthy tissues, complicated placement and removal procedures associated with its application seriously constrained implementation of this therapy regimen. The use of biodegradable devices which are pre-loaded with radioisotopes, thus eliminating the need for their post-treatment extraction, was envisioned by the inventors of the present application as a mean to remove much of the complexity and discomfort associated with brachytherapy.
The inventors of the present invention have developed a novel radioactive and biodegradable composite which was found suitable for local internal radiation therapy, such as branchytherapy. This composite comprises a polymeric matrix embedded with a hydrophobic organic compound, such as a lipid, associated with at least one beta- or gamma-emitting radio-isotope.
The inventors have found that two general sub-groups of such a composite were suitable for internal radiation therapy:
The composites of the invention posses the following properties:
(i) they are fully biocompatible;
(ii) they are fully biodegradable, thus eliminating the need for post-treatment extraction;
(iii) they are loadable with therapeutic doses of radioactive compounds without substantially affecting the biocompatibility and/or biodegradability of the matrix;
(iv) the composites' degradation rate and residence time in the body are controllable and engineerable according to the therapeutic needs;
(v) they are useful in the treatment of diseases or disorders such as neoplasms with minimal effect to neighboring tissues and/or organs;
(vi) they may be used in combination therapies;
(vii) apart from the radioactive compound the composites may be loaded with other agents such as antineoplastic agents; and
(viii) the release of the radioactive compound into neighboring healthy tissue is substantially undetectable.
Thus, in a first aspect of the present invention, there is provided a composite of a polymeric matrix embedded with at least one hydrophobic organic compound associated with at least one radioactive atom, wherein said at least one hydrophobic organic compound associated with said at least one radioactive atom is substantially non-leachable from said matrix.
As used herein, the term “composite” refers to the product produced by combining the polymeric matrix and the organic compound employed in the invention. The embodiment of the hydrophobic organic compound in the matrix is preferably continuous and homogenous. As used herein, the term “embedded” or any lingual variation thereof contemplates any manner by which the radioactive compound is incorporated into the polymeric matrix, including for example: attachment to a monomer of the polymer and having the radioactive compound a part of the polymerization; distribution of the radioactive compound throughout the polymeric matrix; entrapment of the radioactive compound in voids within said matrix; random dispersion of the radioactive compound within the polymeric matrix; encapsulation inside the polymeric matrix, etc. The degree and uniformity of the embedment may also be a result of chemical and/or physical interactions between the matrix and the radioactive compound. Such interactions may be acid-base, hydrophibicity-hydrophilicity, ionic interactions, complexation, chelation, etc.
The term “polymeric matrix” refers to a biocompatible and biodegradable polymer, preferably to an organic polymer, in which the radioactive organic compound is embedded. Non-limiting examples of such polymers are polyurethane, polypropylene, polypropylene terephthalate, polyphenylene oxide, polyphenylsulfone, polyether sulfone, phenyletheretherketone, polyetherimide, silicone and liquid crystal polymer. Non-limiting examples of biocompatible and biodegradable polymers are polyglycaprone, polyglactin and polydioanone.
The polymeric matrix employed may be formed, as will be discussed hereinbelow, from the monomers of the polymer (by polymerization), from the polymer itself under appropriate conditions (such as those leading to hydrogelation) or in the presence of other agents such as curing or crosslinking agents which may be necessary in order to achieve the formation of a desired matrix.
In some embodiments, the polymer is a curable or crosslinkable polymer. The term “crosslinking”, as may be known to a person skilled in the art, refers to the reaction between a polymer and a second chemical entity which may be another polymer or a chemical compound to which the polymer is reactive. The crosslinking may for example be achieved by employing appropriate chemical conditions, as may be known to a person skilled in the art and demonstrated in the Examples. For example, the crosslinking may be of a polymer having repeating units which bare at least one amino group (such as in the case of chitosan) per unit, capable of crosslinking with a dialdehyde-containing compound. The crosslinking may also, for example, be affected by the addition of multivalent cations such as calcium.
The polymer consisting the matrix, in a most preferred embodiment of the invention, is a polysaccharide. The polysaccharide may have the same repeating monosaccharide units (such as in the case of cellulose) or different repeating units (such as in the case of alginic acid), may be natural, synthetic or semisynthetic (modified), branched or linear.
The polysaccharide may be employed in its salt form, namely charged. Typically, the salts are of the polysaccharide and cations such as calcium, magnesium, barium, aluminum, ferrous and ferric, and others.
Non-limiting examples of polysaccharides are alginic acid, amylopectin, amylose, arabinoxylan, cellulose, chitin, chitosan, chondroitin, galactoglucomannan, glucomannan, glycogen, guar gum, heparin, hyaluronic acid, inulin, pectin, and xyloglucan.
Non-limiting examples of crosslinking agents are glutaraldehyde, diaminododecane, and divinyl glycole. Crosslinking may also be achieved between two different polysaccarides.
In one embodiment, the polysaccharide is a curable or crosslinkable polysaccharide having monosaccharides which have at least one amino group (such as in the case of chitosan), which can be crosslinked with dialdehyde-containing crosslinking compounds (such as glutaraldehyde).
The term “hydrophobic organic compound” refers to an organic compound, i.e. having a carbon basis, which is substantially insoluble in water and which may be fully synthetic, partially synthetic or natural. The organic compound is additionally one which is biocompatible and which does not exert any additional toxic effect, apart from the effect exerted by the radiation, on the tissue or organ in which the composite is implanted. Within the context of the present invention, the hydrophibicity of the organic compound should substantially remain the same regardless of the type of association between the organic compound and the radioactive atom. The association between the organic compound and the radioactive atom may depend on the nature of each of the partners.
In some embodiments, the association is via at least one chemical bond, namely the two may be held together via bonding such as covalent, ionic, hydrogen, van der waals, coordination, etc. Preferably, in such embodiments, the association is of a covalent or coordinative nature.
In other embodiments, the association between the hydrophobic organic compound and the radioactive atom is physical, namely the radioactive compound may be entrapped or encapsulated such as in the case of micellar systems and inclusion complexes, e.g. liposomes, dendrimers, etc.
Preferably, the hydrophobic organic compound employed in the composite of the invention is selected from lipids such as cholesterol and norcholesterol, fats such as triglycerides, hydrocarbons of various chain lengths, and others.
The term “radioactive compound” refers herein to the hydrophobic organic compound when associated with at least one radioactive atom. The “radioactive atom” is preferably selected from: (a) gamma emitting radio-isotopes; (b) beta emitting radio-isotopes; or (c) a combination of a gamma emitting radio-isotope and a beta emitting radio-isotope.
In one embodiment, the radioactive atom is selected from radioactive halogens. In this case, the radioactive halogens (i.e. Br, Cl, I and F) are chemically bonded to the organic compound via covalent bonding.
The hydrophobic organic compound associated with the radioactive atom is said of being “substantially non-leachable from the matrix”. This means that regardless of the type of association between the organic compound and the radioactive atom (e.g. covalent, ionic, etc), the radioactive atom does not dissociate from the hydrophobic organic compound. Preferably, the radioactive compound (as defined above) as a whole does not dissociate from the matrix in which it is entrapped until sufficient disintegration of the matrix allows such dissociation. As discussed above, the radioactive compound substantially does not leach from the matrix. However, with the dissociation of the matrix, some leaching may occur. Preferably, the dissociation from the disintegrated matrix occurs after the radiation has substantially decayed.
In a preferred embodiment of the invention, there is provided a composite of a polymeric matrix, preferably a polysaccharide matrix, more preferably a hydrogel thereof, embedded with at least one hydrophobic organic compound, covalently bonded to at least one radioactive halogen, wherein said organic compound covalently bonded to said radioactive halogen is substantially non-leachable from said matrix. Preferably, the at least one radioactive halogen is selected from radioisotopes of iodine (such as 124I, 125I, 131I) and fluorine (such as 18F). Preferably the radioactive halogen is iodine with the most preferred isotope being 131I.
In another embodiment of the invention, the radioactive atom is selected from beta- and/or gamma-emitting radioisotopes such as 124I, 125I, 127I, 131I, 18F, 90Y, 166Ho, 186Re, 188Re, 90Sr, 226Ra, 137CS, 60Co, 192Ir, 103Pd, 198Au, 99Tc, 201Th, 67Ga, 111In and 106Ru and the organic compound with which said radioactive atom is associated is selected amongst lipids, fats, and hydrocarbons.
Preferably, said lipid is one of cholesterol and norcholeterol.
Thus, in another preferred embodiment of the invention, there is provided a composite of a polymeric matrix, preferably a polysaccharide, more preferably a hydrogel thereof, embedded with at least one lipid associated with at least one radioactive atom selected from beta- and/or gamma-emitting radioisotopes, wherein the association between said at least one radioactive atom and at least one organic compound is via an association selected from ionic bonding, coordination bonding, and intermolecular bonding. Preferably, the association is via coordinative bonding. The beta- or gamma-emitting radioisotope may for example be selected from 124I, 125I, 131I, 90Y, 166Ho, 186Re, 188Re, 90Sr, 226Ra, 137Cs, 60Co, 192Ir, 103Pd, 198Au, 99Tc and 106Ru.
As used herein, the term “association” or any lingual variation thereof, in the context of such an expression as “association between the organic compound and the radioactive atom” refers to the chemical or physical force which holds the two entities together. Such force may be any type of chemical or physical bonding interaction known to a person skilled in the art. Non-limiting examples of such association interactions are ionic bonding, covalent bonding, coordination bonding, complexation, hydrogen bonding, van der Waals bonding, hydrophobicity-hydrophilicity interactions, etc. In some preferred embodiments, the association is via covalent bonding. In other preferred embodiments, the association is via coordinative bonding.
It should be understood to a person skilled in the art that in some cases the associative interactions between two atoms or two chemical entities may involve more than one type of chemical and/or physical interactions.
The present invention thus provides composites of at least one polysaccharide matrix being preferably a hydrogel and an organic compound covalently bonded to at least one radioactive halogen; or with a lipid associated with at least one beta- or gamma-emitting radioactive atom.
In one embodiment, the polymeric matrix is embedded with a single radioactive compound. In another embodiment, the matrix is embedded with two or more different radioactive compounds, which may be of the same organic backbone bonded to different radioactive atoms or isotopes, or may be of different organic structures. For example, in one case the matrix may be embedded with a beta-emitting-cholesterol and gamma-emitting-cholesterol, and in another case be embedded with beta-emmiting-cholesterol and beta-emmiting-norcholesterol.
The composite of the invention may be used to treat a specific localized area (locoregion) of the body of the patient. The composite is fabricated so that it retains the radioactive organic compound for a pre-defined period of time without the radioactive compound substantially leaking or disassociating from the polymeric matrix in which it is embedded.
The polymeric matrix embedding the radioactive organic compound is biocompatible. The terms “biocompatible”, “biocompatibility” or any lingual variation thereof, when used in relation to polymers are art-recognized. For example, biocompatible polymers include polymers that are neither themselves toxic to the host (e.g., an animal or human), nor degrade (if the polymer degrades) at a rate that produces monomeric or oligomeric subunits or other byproducts at toxic concentrations in the host.
In certain embodiments of the present invention, biodegradation generally involves degradation of the polymer in a tissue or a body fluid, e.g., into its monomeric subunits, which may be known to be effectively non-toxic. Intermediate oligomeric products resulting from such degradation may have different toxicological properties, or biodegradation may involve oxidation or other biochemical reactions that generate molecules other than monomeric subunits of the polymer. Consequently, toxicology of a biodegradable polymer intended for in vivo use, such as implantation or injection into a patient, may be determined after one or more toxicity analyses.
It is not necessary that any subject composite have a purity of 100% to be deemed biocompatible; indeed, it is only necessary that the subject composites be biocompatible as set forth above. Hence, the composite of the invention may comprise polymers which are only 99%, 98%, 97%, 96%, 95%, 90%, 85%, 80%, 75% or even less biocompatible, e.g., including polymers and other materials and excipients described herein, and still be biocompatible.
To determine whether a polymer or other material is biocompatible, it may be necessary to conduct a toxicity analysis. Such assays are well known in the art, and are exemplified hereinbelow. In addition, polymers and composites of the present invention may also be evaluated by well-known in vivo tests, such as subcutaneous implantations in rats to confirm that they do not cause significant levels of irritation or inflammation at the subcutaneous implantation sites.
In the preferred embodiments, the polymeric matrix is adapted to be degraded and/or absorbed by the body. In such cases, the polymer is eliminated by the body over time, and the dissolution time is preferably chosen to be sufficiently greater than the radioactive half-life of the radioactive material, insuring that the remaining radioactivity does not poses any hazard to body tissues as it migrates from the treatment volume or site. As such, the polymeric matrix is based on biocompatible materials such as the polysaccharides mentioned hereinbefore, which are biodegradable and evoke no toxic response when released into the body.
The term biodegradable is art-recognized, and includes polymers, composites and formulations comprising thereof, such as those described herein, that are intended to degrade during in vivo use, such as implantation. In general, degradation attributable to biodegradability involves the degradation of a biodegradable polymer into its component subunits, or digestion, e.g., by a biochemical process carried out for example by enzymes, of the polymer into smaller, non-polymeric subunits.
Two different types of biodegradation may generally be identified. For example, one type of biodegradation may involve cleavage of bonds in the polymer matrix. In such biodegradation, monomers and oligomers typically result, and even more typically, such biodegradation occurs by cleavage of a bond connecting one or more of subunits of a polymer. In contrast, another type of biodegradation may involve cleavage of a bond internal to side chain or that connects a side chain to the polymer backbone. As used herein, biodegradation encompasses both general types of biodegradation.
The degradation rate of a biodegradable polymer often depends in part on a variety of factors, including the chemical identity of the linkage responsible for any degradation, the molecular weight, crystallinity, biostability, and degree of cross-linking of such polymer, the physical characteristics of the implant (for example porosity), shape and size, and the mode and location of implantation. For example, the greater the molecular weight, the higher the degree of crystallinity, and/or the greater the biostability, the slower the biodegradation.
As stated herein, the composite of the invention is preferably a hydrogel of a polysaccharide. The term “hydrogel” is art-recognized and typically refers to a colloidal system with at least two phases, in which a dispersed phase (being in this case the polymer or polysaccharide) coexists with a continuous phase (being typically water) to form a continuous three-dimensional network which is generally viscous and jellylike.
The hydrogel may be prepared by a number of methods known in the art. One such exemplary method involves the admixing of a suitable polymer or a polysaccharide and the radioactive hydrophobic organic compound in pure water or in an aqueous solution containing e.g. acid. The hydrogelation may be spontaneous or may be achieved by e.g. heating the solution to a specific temperature or adjusting the pH.
Additionally, in order to achieve specific or improved physical matrix, a crosslinking agent may be added.
In order to achieve controlled or predetermined biodegradability, the hydrogel may be further treated, e.g. by washing or incubation, in appropriate conditions as discussed hereinbelow.
The composite of the invention may also be fabricated as a therapeutic source. As used herein, the term “therapeutic source” refers to a device that can be fabricated from the composite of the invention. This source may take on any structure or shape as may be determined suitable by the medical practitioner. The source may be any piece of the composite, or a well structured source such as in the shape of a cylindrical rod, a hallow rod, a suture, a mesh, a film, a sheet or spheroidal. While preferably the composite of the invention is biodegradable and requires no post-treatment extraction, the therapeutic source, as defined herein, may be designed and used as an extractable temporary implant or as a permanent implant.
Preferably, the composite or source fabricated therefrom or a composition comprising thereof is suitable for internal local radiation therapies such as brachytherapy, as well as for intracavity, interstitial, intraluminal and intravascular radiation therapy and/or for injection directly into a tumor.
The composite of the invention may also be contained within a second layer of polymeric material. This type of second polymeric material is typically not radioactive. The second polymeric material may in other cases be the core around which the composite of the invention is fabricated.
In certain embodiments, other materials may be dispersed in the polymer matrix. Such materials may for example be antineoplastic agents used for combination therapy of the same neoplasm or a different one, to alter the physical and chemical properties of the resulting polymer, including for example, the release profile of the resulting polymer matrix, etc. Examples of such materials include active ingredients, biocompatible plasticizers, delivery agents, fillers and the like.
The composite may also be suitable for the manufacture of or use as a radioactive medical implant for placement at a surgical site or a body cavity, i.e. any space or void within the body or in one of its organs, which may be normal or pathological. The medical implant may also serve as a supportive structure for preventing tissue collapse and for reducing tissue deformity.
In another aspect of the present invention, there is provided a method for the preparation of a composite of the invention, said method comprising:
In some embodiments, the method may utilize a pre-formed polymer or polysaccharide. In other embodiments, the method may further comprise the step of polymerization. In these cases, the method may involve the admixing of monomers or short oligomers or pre-polymers with the radioactive hydrophobic organic compound in a suitable media which would allow in situ polymerization and gelation or hydrogelation, thereby affording a composite of the invention.
In other embodiments, the polymer or polysaccharide employed may require curing or crosslinking. Such curing or crosslinking may be chemical, e.g. require the presence of a crosslinking agent, or physical, e.g. radiation curing. One such example of crosslinkable polysaccharide is chitosan which may be crosslinked in the presence of dialdehyde-containing agents.
The method may further also comprise the step of washing or incubating the hydrogel obtained in a suitable media. The degree of polymerization (cross-linking) of the gel and post-curing processes, such as washing and incubation in various media, may be used to control the viscosity of the hydrogel obtained, its elastic/plastic properties, and its in vivo degradation profile. Specifically, a composite having predetermined (controllable) degradation properties may be prepared by incubating the composite of the invention in media with different properties selected from buffer capacity, pH, ionic strength and osmolarity, each capable of modifying the biodegradation properties of the polymer.
For example, when the composite of the invention is subjected to rinsing with a phosphate buffer solution a slow degrading composite (SDC) is obtained; whereas, when water is used as the rinsing medium, a fast degrading composite (FDC) is obtained.
Thus, the present invention further provides a method for controlling the biodegradability of a composite of the invention, said method comprising washing or incubating a composite as disclosed herein with a suitable media. In one case, the suitable media is water, thereby obtaining a fast degrading composite (FDC). In another case, the media is a phosphate buffer solution, thereby obtaining a slow degrading composite (SDC). Preferably, the FDCs of the invention are prepared such that they biodegrade during about a week to about one month from the time of application, whereas the SDCs are prepared such that they biodegrade during a period of about up to three months from the time of application.
In still another aspect of the present invention, there is provided a method for treating a disease in a subject, said method comprising instilling into an anatomic area affected by said disease (being pre- or post-surgery) a therapeutically effective amount of a composite of a polymeric matrix embedded with at least one hydrophobic organic compound associated with at least one radioactive atom, wherein said at least one hydrophobic organic compound associated with at least one radioactive atom is substantially non-leachable from said matrix.
The invention further provides a method for the treatment of a locoregional disease after tumor resection, said method comprising instilling into an anatomic area affected by said tumor a therapeutically effective amount of a composite of a polymeric matrix embedded with at least one hydrophobic organic compound associated with at least one radioactive atom, wherein said at least one hydrophobic organic compound associated with at least one radioactive atom is substantially non-leachable from said matrix.
In one embodiment, the instillation may be in the form of a therapeutic source. The instilled composite or therapeutic source may be for example into a site of a surgically removed tumor or into or around a body organ.
As used herein, the term “treating” or any lingual variation thereof refers to the instillation, i.e. administration of the composite of the invention with the intention of preventing a disease, disorder or a certain condition associated therewith from occurring, inhibiting the disease or arresting or slowing its progression, relieving the symptoms associated with the disease, and ameliorating at least one symptom of the disease.
As used herein, the term “anatomic area” refers to any part of the body, being a tissue or an organ of the body or a cavity. The term “tissue” refers to an aggregation or collection of morphologically similar cells and associated accessory and support cells and intercellular matter, including extracellular matrix material, vascular supply, and fluids. The tissue may be any tissue of the body including blood. The term “organ” refers to any part of the body of an animal or of a human that is capable of performing some specialized physiological function. The term may include any part of such an organ or a collection of one or more of such organs. Non-limiting examples of organs include the heart, lungs, kidney, ureter, urinary bladder, adrenal glands, pituitary gland, skin, prostate, uterus, reproductive organs (e.g., genitalia and accessory organs), liver, gall-bladder, brain, spinal cord, stomach, intestine, appendix, pancreas, lymph nodes, breast, salivary glands, lacrimal glands, eyes, spleen, thymus, bone marrow.
In one embodiment, said disease is a neoplasm, which may be benign or malignant. The neoplastic diseases generally include cancers of the prostate, lung, cervical, colorectal, pancreas, breast, head and neck, melanomas or solid tumors of soft tissues.
Non-limiting examples of neoplastic diseases are adrenocortical carcinoma, anal cancer, bladder cancer, brain tumor, brain stem glioma, cerebellar astrocytoma, cerebral astrocytoma, ependymoma, medulloblastoma, supratentorial primitive neuroectodermal and Pineal tumors, visual pathway and hypothalamic glioma, breast cancer, carcinoid tumor of the gastrointestinal, cervical cancer, colon cancer, endometrial cancer, esophageal cancer, extrahepatic bile duct cancer, Ewing's family of tumors, extracranial germ cell tumor, eye cancer, intraocular melanoma, gallbladder cancer, gastric cancer, germ cell tumor, gestational trophoblastic tumor, head and neck cancer, hypopharyngeal cancer, islet cell carcinoma, laryngeal cancer, lip and oral cavity cancer, liver cancer, lung cancer, malignant mesothelioma, melanoma, merkel cell carcinoma, metastatic squamous neck cancer, plasma cell neoplasms, mycosis, myelodysplastic syndrome, myeloproliferative disorders, nasopharyngeal cancer, neuroblastoma, oropharyngeal cancer, osteosarcoma, ovarian epithelial cancer, ovarian germ cell tumor, ovarian low malignant potential tumor, pancreatic cancer, paranasal sinus cancer, parathyroid cancer, penile cancer, pheochromocytoma cancer, pituitary cancer, prostate cancer, rhabdomyosarcoma, rectal cancer, renal pelvis and ureter, salivary gland cancer, Sezary syndrome, small instetine cancer, soft tissue sarcoma, stomach cancer, testicular cancer, tymoma, thyroid cancer, urethral cancer, uterine cancer, vaginal cancer, vulvar cancer, and Wilm's tumor as well as metastasis of the above.
Preferably, the neoplastic disease is breast cancer, liver cancer or lung cancer, any subtype thereof as well as their metastasis.
When, for example, the breast tumor is excised surgically, the most likely site of recurrence is known to be in the region immediately surrounding the excised tumor. For such a reason, the surgery is typically followed by extension radiation therapy in this region which increases the chance of healthy tissues being damaged by the radiation. Thus, by constructing the composite of the invention into forms such as radioactive sutures, radioactive mesh, etc., a simpler and safer method of irradiation is accomplished.
Methods for treating cancers according to the present invention involve gaining access to the anatomic site where the cancer is to be treated or its evolution prevented and instilling therein the composite of the invention, a composition containing thereof or a source manufactured therefrom. The term “instilling” or any lingual variation thereof, in its broadest scope refers to any type of administration or placement of the composite of the invention in the anatomic site of a subject. Such instillation may employ any method known to a medical practitioner. Typically, access to the anatomic site may be gained by surgical or other invasive procedures, e.g. laparoscopy. Non-surgical methods for the instillation of the composite may also be employed. One such method makes use of colonoscopy for the delivery of the composite of the invention to the colon for the treatment of, e.g. colorectal cancer. In the case, for example, of skin cancers or disorders associated with the skin, the composite of the invention may be placed in direct contact with the sldn, typically without the need of employing any invasive procedures.
Generally, the composite of the invention, a composition comprising thereof or a source manufactured therefrom may be instilled into any anatomic site of the body (human or animal) by accessing the anatomic site and placing the composite of the invention therein. Depending on the method employed for accessing the anatomic site, i.e. surgical methods, non-invasive methods or non-surgical methods, further post operative methods may be employed.
The instillation allows preventing, minimizing, delaying or arresting the occurrence or recurrence of a cancer in a patient who is at risk of developing the disease. The composite, composition or source may be instilled to permanently remain in the anatomic site or degrade over time, resorbed by the tissues and metabolized. Repeated instillations of the composite, composition or source may be undertaken based on the specific formulations.
Combination therapies for advanced cancer patients may also be envisioned and thus fall within the scope of the present invention. Some combination therapies may involve the instillation of a composite of the invention into the anatomic site simultaneously with the administration of another treatment such as systemic chemotherapy, locoreginal radiation therapy, cryotherapy, resection of tumors, and others.
The composite of the invention may also be used in non-neoplastic diseases and conditions such as restenosis. Restenosis is the narrowing of a blood vessel (usually a coronary artery) following the removal or reduction of a previous narrowing (angioplasty). Due to cell proliferation and/or plaque formation occurring in more than 40% of all post-angioplasty procedures, physicians are forced to perform complex and life threatening procedures such as coronary artery heart bypass surgery. The composite of the invention may be utilized in such cases as well as, for example, a stent-like radiation source for the delivery of a dose of radiation to the artery walls. This direct and local radiation assists in reducing the chance of restenosis and the likelihood for more complex and dangerous post-angioplasty procedures.
In order to understand the invention and to see how it may be carried out in practice, preferred embodiments will now be described, by way of non-limiting examples only, with reference to the accompanying drawings, in which:
A person of skill in the art would recognize that the examples provided herein are presented as non-limiting embodiments of the present invention. Thus, for example, a person skilled in the art would be of the knowledge to replace one polysaccharide under another employing the necessary modifications.
Increasing amounts of glutaraldehyde (GA) were used to prepare a series of chitosan (Ct) hydrogels with different crosslinking densities that were characterized by eosin adsorption. Typically, the adsorption process lasted about three hours (
Differential Scanning Calorimetry (DSC) analysis provided information about the thermal changes of the hydrogels, from which the exothermic enthalpy and Tg was computed.
The swelling of G10 gel in increasing ionic strength, osmolarity and pH is depicted in
G10 was further used to produce two types of implants, slow degrading composite (SDC) and fast degrading composite (FDC) of G10, differing from each other by their in vivo degradation rates. The former was obtained by dialysis against PBS and the latter was dialyzed against water, which resulted in different swelling properties. The degradation properties of the two gels were tested following SC and IP implantation in rats. No weight loss of the SDC could be detected over 28 days for both SC and IP implantations. In contrast, only 19.8±9.5 and 9.2±6.5% of the FDC was lest after 14 days of SC and IP implantations, respectively (
Histological observations shown in
The implantation of the 131I-NC loaded implants was shown to delay the progression of solid tumors, as shown in
Kaplan-Meier analysis of mice survival of the different groups showed increased survival of the treated group by 120% compared to control groups (42 and 35 days respectively) (
The long-term survival rate among women who undergo breast-conserving surgery is the same as that among women who undergo radical mastectomy. Moreover, the cumulative incidence of locoregional tumor recurrent in breast was significantly decreased in women who underwent lumpectomy and breast irradiation, as compared with women who underwent lumpectomy without irradiation.
On this background, the effect of the composites of the invention loaded with 131I-NC in the prevention of locoregional recurrence in xenograft breast cancer model was assessed. During surgery to implant the composites of the invention, 4T1 cells (10% of the amount needed for primary tumor induction) were spread subcutaneously in the surgical cavity, to mimic cancer cell spillage in the tumor bed causing the locoregional recurrence. Kaplan-Meier analysis showed that the survival of the group treated with hydrogels loaded with 131I-NC was 69.2%, as compared to death of all mice in the group treated with empty gels and in the untreated group (
Without wishing to be bound by theory, the elimination of the radioactivity from the site of implantation in vivo is a result of two parallel pathways: the radioactivity decay and the biological elimination of the radioactive material. It was earlier shown that the elimination of 131I-NC depended on the degradation of the hydrogel, due to the hydrophobic nature of the 131I-NC and the hydrophilic nature of the hydrogel. The total radioactivity elimination constant was calculated as the slope of linear regression of the natural logarithm of the remaining fraction of radioactivity with time (
In summary, implantation of composites of the invention such as hydrogels loaded with 131I-NC in the vicinity of tumor, reduced tumor progression rate by two weeks compared with the control groups. Implantation of the hydrogels loaded 131I-NC in the tumor bed in a residual disease model prevented tumor recurrence and increased survival by 69% as compared with a complete mortality of the control groups. Histopathological analysis of the tumor bed and distant organs after implantation of the hydrogels showed that the implants of the invention were safe and biocompatible.
A—Biocompatibility Evaluation and Mode of Degradation of Crosslinked Chitosan Hydrogels after Subcutaneous and Intraperitoneal Implantation in Rats
One hundred milligrams of Ct was dissolved in 10 ml of 1M acetic acid (Frutarom, Israel), and heated to 100° C. Glutaraldehyde (GA) solution (25% w/v in water) was then added while stirring. A gel was formed immediately and the stirring was stopped. Assuming complete reaction between one molecule of GA and two glucosamine repeating units, the following Ct:GA molar ratios were examined in different studies: 1:5, 1:7.5, 1:10, 1:12.5, 1:15, 1:17.5 and 1:20. These ratios are hereby denoted as G5, G7.5, G10, G12.5, G15, G17.5 and G20 respectively. Excess of GA was removed by dialysis until no traces of GA could be detected at 235 nm (polymeric GA) and 280 nm (monomeric GA) (Uvikon 930, Kontron Instruments, Switzerland) in the rinsing medium.
The crosslinking density of the gels was quantified by adsorption measurements of the negatively charged dye eosin from a hydroalcoholic solution. In different studies about 0.2 g of each gel was incubated in 2 ml of 0.05 mg of eosin in ethanol:water 1:1 solution for 10, 30, 60 and 180 minutes at room temperature. The gels were then removed and the eosin concentration in the incubation medium was measured spectrophotometrically (520 nm), using a six-point calibration curve. The gels were then rinsed with water, dried in acetone (48 hours) and weighed. The amount of eosin adsorbed, which was calculated from the initial and final concentrations in the bathing solution, was normalized to the dry weight of each gel.
The gels' crosslinking density was also characterized by differential scanning calorimetry (DSC) analysis. The change in heat capacity of the pre-dried gels was measured in a temperature range of 25-175° C., at a rate of 10° C./min, under N2 flow of 1 ml/min (Mettler Instruments, Switzerland, TA 4000, equipped with TC II TA processor). DSC curves were plotted and the glass transition temperature (Tg) and enthalpy (ΔH) were computed by the apparatus program.
Typically, the adsorption process lasted about three hours (
DSC analysis provided information about the thermal changes of the hydrogels, from which the exothermic enthalpy and Tg was computed.
Cubic (S-4 mm) specimens from each gel were cut by a scalpel and tested in a texture analyzer (TAXT Plus, Texture Technologies, USA), at a rate of 0.05 mm×sec−1 (compression). Young's modulus of elasticity (E) was calculated using to the following equation:
E=(F/A)/(ΔL/L0) Eq. 1
where F is the tensile force (in gF), A is the cross section area of the specimen (cm2), ΔL is the specimen's strained length (mm) and L0 is the initial length of the gel's specimen (mm).
The Young Modulus was calculated from the elasticity studies (
Since the gels are designed to be implanted in the hydrated form, the effect of ionic strength, osmolarity and pH on their swelling properties were studied in vitro. In separate studies specimens of G10 were incubated in increasing concentrations (0, 10, 50, 100, 150, 200 or 400 mM) of NaCl (increasing ionic strength) or glucose (increasing osmolarity). Similarly, the G10 specimens were incubated in 10mM phosphate buffer at different pH values (3, 5, 6, 7, 8 and 10). The incubation lasted 12 hours, after which time the specimens were rinsed with water, weighed (WS), dried, weighed again (WD) and the WS/WD ratio was calculated.
The swelling of the G10 gel in increasing ionic strength, osmolarity and pH (expressed as the ratio of its weight in a swollen state to its dry weight) is depicted in
Two formulations of G10 were prepared and each was subjected to a different mode of dialysis after crosslinking. The slow degrading G10 (SDC) was prepared by dialysis against PBS (1 mM, pH=7.4), while the fast degrading G10 (FDC) was prepared by dialysis against water.
In separate studies the SDC and FDC were implanted both intraperitonealy (IP) and subcutaneously (SC) in the rat. The former was conducted by laparotomy through a midline incision in the anesthetized rat, and placement of the gel specimen between the intestine and the peritoneum, approximately 1 cm left to the incision. The latter was conducted by retracting both muscles and skin to form a cavity into which the gel specimen was inserted, approximately 1 cm left to the incision. The abdominal cavity and skin were then closed using a 3-0 vicril running suture (Johnson & Johnson Medical). Following surgery the rats were supervised until complete recovery and then normal diet was resumed. At 0, 1, 3, 7 and 14 days, for the FDC, and 0, 3, 7, 14 and 28 days for SDC, four rats from each group were sacrificed. The gels were retrieved, rinsed in water, dried and weighed (WRem).
The extent of the in vivo degradation (as % of initial amount) was assessed from the ratio of the gels' dry weight before and after implantation (the gels were implanted in the hydrated form and lost water during the course of the rat studies). The weight ratios hydrated/dry of the gels prior to implantation were measured and found to be 52.0±0.9 and 198.1±1.9 for the SDC and FDC, respectively. W0 was calculated from the above ratios. The extent of gel degradation, in percent of initial amount (% Remained), was calculated using the following equation:
% Remained=(WRem/W0)×100 Eq. 2
where, WRem is the dry weight of the gel debris retrieved at the end of each implantation study and W0 is the initial dry weight of the respective gel.
G10 was further used to produce two types of implants, SDC and FDC, differing from each other by their in vivo degradation rates. The former was obtained by dialysis against PBS and the latter was dialyzed against water, which resulted in different swelling properties (wet/dry weight ratios of 198.1±1.9 compared with 52.0±0.9, respectively). The degradation properties of the two gels were tested following SC and IP implantation in the rat. No weight loss of the SDC could be detected over 28 days for both SC and IP implantation. In contrast, only 19.8±9.5 and 9.2% 6.5% of the FDC was left after 14 days of SC and IP implantation, respectively (
The hydrophobic dye Sudan black (SB) was loaded into the SDC and FDC gels to allow further insight into the degradation kinetics of the two types of implants in vivo. SB was dispersed in the acidic Ct solution to obtain a final SB:CT ratio of 1:100. After crosslinking with GA to obtain an SB loaded G10 hydrogel, SB loaded SDC and FDC gels were prepared by dialysis, as described above. After weighing, the two products were implanted SC and IP in the anesthetized rats. The rats recovered and were maintained with free access to normal rat chow and water. At 0, 1, 3, 7 and 14 days, for the FDC implanted group, and 0, 3, 7, 14 and 28 days for SDC implanted group, four rats were sacrificed at each time point. The gels or gel debris were located, separated from the tissues, rinsed and soaked in acetone for 48 h in a sealed beaker. The concentration of the extracted SB (SBRem) was measured spectrophotometrically at 600 nm and the fraction of SB released during implantation was calculated. The initial amount of SB in the gels (SB0) was determined by acetone extraction at time 0. The amount of SB remaining in each gel specimen removed from the rats at each time point was measured similarly and normalized to the gel weight. SB0 was calculated from the weight of the implanted gel and the calculated SB/Gel ratio, while the fraction of SB released from the gel was calculated as follows:
% SB Released=(SB0−SBRem)/SB0×100
To verify the degradation results and to investigate the gels' ability to serve as platforms for hydrophobic probes, both types of implants were loaded with SB. The study of the SB release kinetics in vivo revealed that only 13.6±8.3 and 18.7±1.4% SB were released from the SDC after 28 days of SC and IP implantation respectively. However, almost complete SB release occurred during the first week of SC and IP implantation of FDC, indicating that its degradation was the cause for the accelerated release of the dye (
0.2 ml of 131I-NC (0.2 mCi) (CIS Bio International, France) was dispersed in 10 ml of Ct solution (1% in acetic acid 1M), heated to 100° C. and 1.2 ml glutaraldehyde solution (25% w/w) was added to form a gel. Gels were dialyzed against PBS pH 7.4 (1 mM) for 24 hours to obtain the SDC gels, and 0.5 g specimens of the gels were implanted in the left pectoral region of the rat. Scintigraphy was performed at 0, 3, 13 and 30 days after implantation. Each rat was imaged for 15 min under anesthesia, using a helix dual-head camera (Elscint, Haifa, Israel) and a high-energy, high-resolution collimator. Data was analyzed on a Xeleris program (GE Healthcare), and regions of interest were drawn on each focus. The total number of counts in each region was calculated and percent of activity in each regions of interest was calculated as follows:
% Activity=(At/2−t/8.02)/A0×100 Eq. 4
where t=time (days), At=counts at t, A0=counts at t0, and 8.02 days is the half-life of 131I.
The effect of the two types of implants, FDC and SDC on the surrounding tissues was examined histologically. Tissue specimens taken at the site of implantation were collected at 14 days for FDC and 28 days for SDC. All specimens included debris of the implant itself.
The specimens were fixed in 4% buffered formaldehyde, dehydrated, embedded in paraffin blocks and four-micron slices were sectioned and stained with hematoxylin and eosin. The sections were examined microscopically for a possible inflammatory response and evaluation of the thickness of the peri-implant fibrotic capsule. Tissues specimens containing debris of Vicryl™ biodegradable sutures served as controls for foreign body tissue interactions and were collected after 28 days.
Histological observations shown in
In separate studies SDC and FDC (1 g/Kg body weight) were implanted each intraperitonealy (IP) and subcutaneously (SC) in the rat. IP implantation was conducted by laparotomy and placement of the gel approximately 1 cm left to the midline, in the peritoneal cavity. SC implantation was conducted by retracting the muscles and skin, and mounting the gel approximately 1 cm left to the midline, between the muscles and the skin. After the gels' implantation the abdominal cavity and skin were sutured and the rats were allowed to recover. At 0, 1, 3, 7 and 14 days, for the FDG implanted group, and 0, 3, 7, 14 and 28 days for SDG implanted group 4 rats were sacrificed and tissue specimens from the tissues surrounding the implants were rinsed with PBS, fixated with 4% formaldehyde in PBS, dehydrated, embedded in paraffin blocks, sectioned (4 μm) and stained with hematoxylin-eosin for histological examination of tissue reaction.
A minimum of three serial sections of each block was examined microscopically in search of cellular inflammatory response and for measurement of peri-implant fibrotic capsule thickness. The extent of inflammatory response was quantified by assessing the presence of inflammatory cells (polymorphonuclears, lymphocytes, macrophages and foreign body giant cells), fibrin, exudate, necrosis and vascularization. The presence of the above inflammatory markers was scored from (−)=absence, to (+++)=profound presence (Table 1). The peri-implant fibrotic capsule thickness was defined as the distance between the border of the fibrotic tissue, adjacent to the implant, and the muscle or fat tissue adjacent to the fibrotic capsule at the other end. The draining regional lymph nodes were microscopically examined in a search for non-typical reactive response (lymphocytes or macrophages infiltration).
After implanting the two types of gels in two different locations in the rat, inflammation was observed in the tissues surrounding the implants. The variable degrees of inflammatory component infiltration and fibrous capsule formation are summarized in Table 1, which also shows that neither hemorrhage nor necrosis occurred around the implants.
One day after implantation, the inflammatory response in the surrounding tissues was characterized by the typical appearance of polymorpho nuclear cells (PMN) (
The surrounding tissue response to intraperitoneal FDC implants showed a similar pattern to that observed after subcutaneous implants, including some degree of degradation of the implant substance. However, the intraperitoneal tissue response was milder and more indolent creating a thinner capsule measuring in average of 80 μm (
On day 3, the peri-implant tissue response was dominated by activated fibroblasts with mixed cellular response including lymphocytes, macrophages and new vascularization (
The surrounding tissue response to intraperitoneal SDC implants showed a similar pattern to that observed following the implantation of subcutaneous implants. However, the intraperitoneal tissue response was milder and more indolent creating a thinner capsule measuring in average of 80 μm (
In separate studies three doses (1, 5 and 15 g/kg) of SDC gel and 2 cm of 3/0 polyglycolic-polylactic absorbable suture (Vicryl®, Ethicon, Piscataway, N.J., USA) were implanted SC in the back of four groups of rats for the purpose of pathology comparison to a non-treated group (n=15 rats in each group). Five rats from each group were sacrificed at 4, 14 and 30 days and specimens were taken from brain, lung, kidney, liver, spleen and sternal bone marrow for histological assessment of possible tissue injury and presence of microscopic debris of the implanted objects.
Neither the presence of gel fragments nor tissue damage could be observed in the brain, heart, lung, kidney, liver, spleen and sternal bone marrow of the tested rats at all time points (0, 4, 14 and 30 days), after either implantation of the three doses of the gel (1, 5 and 15 g/Kg), or the polyglycolic-polylactic absorbable suture material.
Cubic FDC specimens (s=4 mm) were incubated in different concentrations (0, 1, 5 and 10 mM) of KMnO4 for 3 min. The gels were retrieved, washed twice with water, and incubated separately in 1 ml of hematoxylin (0.05 mg/ml solution) or eosin (0.5 mg/ml solution) for 4 h at room temperature. The concentration of the remaining dye in the incubation medium was measured at 560 nm (hematoxylin) and 520 nm (eosin) and the fraction (percent from initial amount) of dye adsorbed onto the gels was calculated.
FDC gels implanted intraperitonealy in the rat showed signs of partial and total degradation after 7 and 14 days, respectively (
Specimens (0.5 g) of the SDC hydrogels loaded with 131I-NC (see above) were implanted in the left pectoral region of three anesthetized rat, and the tissue response in the peri-implant tissue was evaluated after 30 days, as described above.
In contrast to the mild tissue response to the non-radioactive SDC and FDC implants, SDC loaded with 131I-NC caused a profound inflammation in the tissues surrounding the radioactive gel. In some cases liquefactive necrosis was observed (
B—Tumor Recurrence Prevention by Brachytherapy Using Biodegradable Crosslinked Chitosan Hydrogel Implants Loaded with 131I-Nor-Cholesterol
4T1 cells, from metastatic mouse breast cancer, were cultured at 37° C. in a humidified atmosphere of 5% CO2/air in Dulbecco's Modified Eagle's Medium supplemented with 10% heat-inactivated fetal bovine serum, penicillin G (60 mg/liter), and streptomycin (100 mg/liter). Cells were harvested with Trypsin-EDTA, washed with PBS, and concentrated to 2.5×105 and 2.5×103 cells/ml in PBS for tumor progression and micro-residual disease studies, respectively.
Female, 7-9 weeks, BALB/c Mice were used in this study, which was conducted in accord with the Principles of Laboratory Animal Care (NIH Publication #85-23, 1985 Revision). Anesthesia was performed by an intraperitoneal injection of 100 mg/kg body weight of ketamine (Ketase™, 0.1 g/ml Fort Dodge, USA). Euthanasia of the anesthetized rats was carried out by chest wall puncturing.
A suspension of 4T1 cells (0.2 ml) was subcutaneously injected in the back of sixty female BALB/c mice (5×105 cells/mouse). Mice were observed for a further two weeks for tumor progression. Mice were anesthetized, a 1-cm incision was performed in the back skin adjacent to the tumor, hydrogels were implanted in the vicinity of the tumor and the skin was sutured. The study was divided into three groups, each group contained 20 mice:
group-1 was a non-treated control group in which no hydrogels were implanted;
group-2 was implanted with 0.5 g of empty hydrogel as a vehicle control, and
group-3 was implanted with 0.5 g of 131I-NC loaded hydrogels.
Three mice were sacrificed from each group at 2, 3 and 4 weeks. Tumor and internal organs (lung, heart, liver and kidney) were dissected, weighed and analyzed histologically. Moreover, Kaplan-Meier survival analysis was performed over 6 weeks.
Tumor progression rate in untreated and empty-hydrogels treated groups was 0.11 g/day in the first 21 days after the beginning of the treatment, and no significant progression was detected beyond this time (
In the untreated group and empty-hydrogel treated group mortality initiated at day 17 and completed by day 35 after hydrogel implantation (
Tumor cells (10% of the amount utilized for primary tumor implantation) were spread during the implantation of hydrogels in healthy mice, mimicking micro-residual disease in the tumor bed after surgical tumor removal, and tumor cell spillage during the surgical procedure. Sixty mice were anesthetized, a 1-cm incision was performed in the back skin of the mice, hydrogels were implanted, 0.2 ml of cell suspension (5×103 cells/mouse) were spread dropwise in the implantation site and the skin was sutured. The study was divided into three groups as described above. Histological analysis of organs (lung, heart, liver, kidney and tissue at the site of cell implantation) was performed at 11 weeks after cell injection, and Kaplan-Meier survival analysis was performed over 20 weeks.
All the animals in the untreated group and empty-hydrogel treated group developed tumors and died after 77 and 84 days after hydrogel implantation, respectively. Only 31% of the group treated with 131I-NC loaded hydrogels developed tumors and died 77 after hydrogel implantation. However, 69% of this group did not develop any tumors and continued to survive until the study was stopped after 160 days (
Specimens from tumor bed and distant organs (lungs, heart, liver spleen) were dissected from mice at 14 days for the tumor progression model and 80 days for the tumor recurrence model. Specimens were then rinsed with PBS, fixated with 4% formaldehyde in PBS, dehydrated, embedded in paraffin blocks, sectioned (4 μm) and stained with hematoxylin-eosin for histological examination of tumor progression or recurrence and metastasis in distant organs (
Specimens (0.5 g) of 131I-NC loaded hydrogels were implanted subcutaneously in the back of four mice. Scintigraphy was performed at 0, 4, 14 and 30 days after implantation. Each mouse was imaged for 10 min under anesthesia, using a helix dual-head camera (Elscint, Haifa, Israel) and a high-energy, high-resolution collimator. Data was analyzed on a Xeleris program (GE Healthcare), regions of interest were drawn on each focus, and the total number of counts in each region was obtained.
Data from the imaging experiments was expressed as natural logarithm of the fraction of the radioactivity remaining with time after implantation, depicted as open circles in
Eq. %: Q=Q0e−λt,
which was developed to Eq. 6: −Ln(Q/Q0)=λt.
Linear regression was performed on these results to obtain the elimination constant (λ=0.136 Day−1). The elimination constant (λ) is composed of the radioactive decay constant (λR) and the biological elimination constant of the isotope (λB) as described in Eq. 7. The discontinuous line in
| Filing Document | Filing Date | Country | Kind | 371c Date |
|---|---|---|---|---|
| PCT/IL06/01259 | 11/1/2006 | WO | 00 | 1/5/2009 |
| Number | Date | Country | |
|---|---|---|---|
| 60731872 | Nov 2005 | US |
