BACKGROUND OF THE INVENTION Atherosclerosis and stenosis can be interventionally treated by catheter- assisted balloon angioplasty or atherectomy. However, these procedures can cause intimal dissection of the tissues, and result in subsequent restenosis due to intimal hyperplasia. The excessive tissue growth on the internal walls of the blood vessel characterizing intimal hyperplasia can be reduced by the use of intra-arterial stents, left in place temporarily or permanently at the site of the angioplasty or atherectomy.

More recently, the use of intra-arterial stents to prevent restenosis has been enhanced by coating the stents with anti-thrombogenic agents, such as heparin. Additionally, beta-and gamma-emitting isotope stent coatings have been shown to be useful in decreasing intimal hyperplasia. However, such radioisotopes have undesirable side-effects.

Investigations of beta radiation emitting stents have shown that the sharp decay of beta radiation may have proliferative effects at the end of the stents.

Therefore, in a low activity range, proliferation is enhanced instead of prohibited.

The main disadvantage of gamma radiation emitting stents is the difficulty of handling of such hazardous radioisotopes in the cardiac catheterization laboratory. There is no standing U. S. regulatory approval for cardiologists to handle gamma rays.

What is needed is an alternative form of radiation for use in surgical devices, such as arterial stents, that does not have the disadvantages associated with prior irradiation sources.

SUMMARY OF THE INVENTION The present invention provides internal surgical devices formed from an X-ray emitting radioactive material. In particular, the present invention provides stents for use in bodily vesicles, such as arteries, which emit X-radiation to reduce the proliferation of cells in close proximity thereto.

Therefore, it is an object of the invention to provide improved radiation emitting surgical devices, including intra-arterial stents.

BRIEF DESCRIPTION OF THE DRAWNGS Fig. 1 shows the decay characteristics of palladium 103.

Fig. 2 shows the decay characteristics of palladium 103 in logarithmic scale.

DETAILED DESCRIPTION OF THE INVENTION The present invention provides that X-radiation can be effectively used in an internal surgical device to inhibit undesired cellular growth. The X-ray emitting source can be located on at least a portion of a stent, wire, catheter or other suitable internal surgical device delivery means. The X-ray emitting source is preferably palladium (Pd 103). Other sources of X-radiation are cesium (Cs 137), or cobalt (Co 57), for example. The invention preferably provides isotopes emitting X-radiation in the range of about 15 to 35 keV.

Palladium 103 is a preferred X-ray source for the present invention. The emission, energy, half-life and penetration profile of palladium 103 is described in Table 1 below, as compared to other radiation sources.

Table 1 Comparison of Various Radiation Sources IsotopeEmissionMax (3-EnergyHalf-Life Penetration Phosphorus 32 e-100% days 3.7 mm Y 90 e-99. 999% 2.3 MeV 64.1 hours 6.0 mm y 0.00095% 0.909 MeV Stronium 89 e-99.999% 1.5 MeV 50.5 days Stronium 90/e-100% 0.5 MeV/28.5 years 0.2 mm y 90 2.3 MeV Palladium 103 X-ray 100% 20.2 keV 16.96 days 10 mm Penetration in Table 1 is determined by the point at which 75% of the radiation stops within living tissue. The decay characteristics of palladium 103 are detailed in Table 2 below.

Table 2 Decay Characteristics of Palladium 103

Isotope Deca Type γ-rays and e- X-rays palladium 100% electron 20.2 keV 103 converson Rh 103 0.023% 357 keV y (daughter) 0.004% 497 keV y 0.04% 39.8 keV y 99.93% <20 keV e- The very low energy electrons are produced by conversion. Their range in tissue is less than 2.3 um and in palladium is less than 0.19 um. Because of their high intensity and the low range, these electrons produce a very localized dose rate at the surface of the stent. The additional very short dose rate can be partially or completely suppressed by additional thin surface coatings, such as with non-radioactive palladium.

Radioactivity is a measure of emitted ionizing radiation (e. g. electrons, X- rays, gamma rays, etc.), measured in counts per second, equivalent to one becquerel ("Bq"). The Curie unit of activity is equal to 3.7 x 101° Bq. The radiation absorbed dose measures the absorbed energy/unit mass, measured in rad or Gray (Gy), wherein one rad equals 0.01 Gray equals 6.2 x 107 MeV/g.

Today, the worldwide equivalent dose rate is used in units Sievert/sec (Sv/s), wherein 1 Sv equals 100 rem.

Typical prior art stents using phosphorus 32 have an activity of about 5 uCi. Using a similar activation for a palladium 103 coated stent results in the following irradiation characteristics.

Table 3 Irradiation Characteristics of Palladium 103 Coated Stents RadiationEner Radiation Activit 20.2 keV X-rays 185,000 Bq 357 keV y-rays 43 B 497 keV y-rays 7 Bq 39.8 keV y-rays 74 Bq <20 keV e-0 to 184,871 Bq (adjustable)

The decay characteristics of palladium 103 are shown in Fig. 1 and in logarithmic scale in Fig. 2. The radioactive decay is not as sharp for X-rays as it is for beta-particles. X-ray is, like gamma radiation, an electromagnetic wave.

Because the attenuation of these waves is exponential, X-ray is also present at a greater distance as compared to beta-radiation. The sharp drop in beta particle activity with respect to the distance from the source could be responsible for an increase in proliferation at the ends of the stents. The present invention provides a solution to these problems.

The devices of the present invention emit X-radiation characterized by a penetration depth of from about 0.1 to 50 mm, and preferably about 10 mm.

The devices of the present invention emit X-radiation characterized by an activity of from about 50 uCi to about 1,000 uCi, and preferably about 300 uCi.

As mentioned, X-radiation emitters offer several advantages over prior radiation sources. First, the dose rate as a function of range is exponential, in contrast to e-emitters, and therefore allows a larger irradiation volume. In addition, high local dose rates at the stent surface (0 to 2.3 um) can be delivered by low energy electrons, if desired. This e-irradiation is freely adjustable over orders of magnitude, starting from zero. Systematic variability of long versus short range dose delivery can be achieved. Even for a 5 uCi stent, the y-ray background can be considered as low. For comparison, the natural activity (mainly y-rays) of the human body is typically 110 Bq/kg. Therefore, a laboratory or clinical environment licensed for X-ray handling can use the activated stents of the present invention. Furthermore, the decay time for palladium 103 of approximately 16 to 17 days is similar to phosphorus 32 activated stents at 14.3 days.

The construction of stents is generally well-known. Stents are cylindrical structures, usually formed from metal mesh or coils, and sized for

insertion into a particular bodily vesicle. Stents are commonly expandable for secure placement within the vesicle. Such stents are produced by Forschungszentrum Karlsruhe GMBH (Karlsruhe, Germany). The invention contemplates that any stent design or other internal surgical device design can utilize the X-ray emitting radioisotopes.

Another surgical device to which the present invention can be applied is the catheter-assisted insertion of an X-radiation source coated wire (s) into any bodily vesicles in need of radiation treatment. Such X-ray emitting radioactive wires can be used, for example, to treat myocardial ablation (HOCUM) or can be constructed into a cage for use in a MAZE procedure. Wires emitting X- radiation can be used for insertion into a fallopean tube, biliary duct, prostate gland, urethra, bladder, cerebro-spinal passages, lymphatic system, endocranial spaces, or, sinus passages. The invention contemplates devices for treating any bodily vesicle or orifice that may benefit from X-radiation. Preferably, such wires are provided with about 100 to 300 uCi X-radiation, but the amount will vary depending upon the procedure and condition of the patient, as can easily be determined by those skilled in the field given the present disclosure.

The device base material is preferably a stainless steel. Alternatively, the base can be constructed of other suitable metals such as aluminum, chromium, copper, gold, iron, molybdenum, nickel, nitinol, palladium, platinum, rhodium, silver, tantalum, titanium, tungsend or zinc. Furthermore, the base may be constructed of any suitable plastic, ceramic, or fiber composition. The base may be pre-treated, such as by polishing, to increase its affinity for the X-radiation emmitting source coating.

The surgical devices of the present invention can be made directly from an X-ray emitting metal or alloy. More preferably, the devices can be made of a base material, and the X-ray emitting source can be coated thereon. The coating process may be achieved for example, by a galvanization process in a solution containing the radiation source or by sputtering the radiation source onto the base, both well-known in the art of metalurgy.

Briefly, in the sputtering, or physical vapor deposition, process coating material is dislodged and ejected from the solid surface due to the momentum exchange associated with surface bombardment by energetic particles. First, gas ions are accelerated by a high voltage, producing a glow discharge, or plasma.

An X-ray source is bombarded in high vacuum by the gas ions. Atoms from the

target are ejected by momentum transfer and move across the vacuum chamber.

Atoms are then deposited on the base to be coated and thereby form a thin film.

The sputtering process is generally described in Bunshah, Handbook of Deposition Technologies for films and Coating: Science, Technology, and Applications, Second Ed., Noyes Publ. New Jersey 1991. The process of galvanization is described for example in Hamann et al., Electrochemistry, Wiley-VCH, New York, 1998, ISBN 3-527-29096-6. Ion implantation techniques known in the art may also be used confer X-ray emitting properties to the surgical device.

The amount and precise location of the radioactive material can be controlled as desired. For example, it may be preferable to concentrate an increased amount of the X-ray emitting material on the ends of a stent in order to more severely inhibit intimal hyperplasia at those locations. The concentration of radiation emitted from a particular region of the device can be increased by increasing the thickness of the radioactive material coating or increasing surface area in the region of interest.

The coating of X-ray radioactive source material can be from about 1 to 10 atom layers. It should be understood that due to the imperfections of the coating techniques, there may be areas of the base that are not entirely coated with the X-ray emitting source. On average, the X-ray emitting source coating is between about 1 nm to 1000 nm, more preferably about 10 nm.

In preferred embodiments, the devices have an additional biocompatible layer on top of the X-ray emitting layer. For example, the biocompatible coating can be hydrophilic polymer such as polyurethane, polymethyl methacrylate (PMMA, or PLEXIGLASTM), polytetrafluouroethylene (PTFE, or TEFLON), polyethylene, polyacrylonitrile, polyamide, polyethylene teraphthalate (PETP, or DACRONTM), polybutylenterepthalate, or polyoxymethylene. Descriptions of various biocompatible polymers can be found in Oberbach, Kunstofftaschenbuch Saechtling, Carl Hanser Verlang Munich, Wein, 1998, ISBN 3-446-19054-6.

The biocompatible coating layer can also be a metal or alloy material.

The biocompatible coating serves to limit any drift or removal of the X-ray emitting source material from the base. The biocompatible coating may contain additional therapeutic agents, such as radiation sensitizers, anti-angiogenic agents, or anti-thrombogenic agents.

The biocompatible coating process may be achieved for example, by a chemical reaction, dipping or sputtering the polymer onto the device by well- known methods. The biocompatible layer can be from about lnm to 1000 nm, preferably about 500 nm. The biocompatible coating is preferably flexible enough to permit expansion of the stent when in position in the vesicle, without causing cracks or fissures in the coating.

In an alternative embodiment, the surgical device is coated with a single layer comprising both a biocompatible material and a X-radiation source. When the device has a single biocompatible material and a X-radiation source layer, it is preferably between about lnm to 1000 nm, preferably about 500 nm thick. In certain embodiments, the polymeric material forms a matrix wherein the X- radiation source, e. g. palladium 103, or other materials can be incorporated into the interstices within the matrix. In certain embodiments, the polymeric material allows the X-radiation source to remain affixed thereto, whereas other materials, such as therapeutic agents, can be delivered from the surgical device when in contact with a patient. For example, therapeutic agents which counteract the side effects of radiation may be incorporated into the coating. In certain embodiments, an additional carbon-based layer, such as graphite, can be disposed on the surgical device, and then overcoated with the biocompatible material layer.

The above-description of the preferred embodiments of the invention is intended to be exempliary of, and not limiting to, the full scope and spirit of the appended claims. It is understood that the invention contemplates many additional modifications, adaptations and alternative embodiments intended to be encompassed by the claims.