Abstract:
A radiotherapy system comprising an X-ray tube operating at 100 to 800 kVp for providing X-ray beams of 30 mm diameter or less and configured to move the entrance beam footprint on the body during irradiation to any arbitrary sequential position set that has been predetermined to limit the intervening tissue dose rate at any one location to a safe level, such that the sum of the skin area traversed during treatment is 20 to 100 times the beam area.

Description:
REFERENCE TO PENDING PRIOR PATENT APPLICATION 
       [0001]    This patent application claims benefit of pending prior U.S. Provisional Patent Application Ser. No. 61/881,274, filed Aug. 23, 2013 by John K. Grady for SMALL BEAM AREA, MID-VOLTAGE RADIOTHERAPY SYSTEM WITH REDUCED SKIN DOSE, REDUCED SCATTER AROUND THE TREATMENT VOLUME, AND IMPROVED OVERALL ACCURACY (Attorney&#39;s Docket No. GRADY-2 PROV), which patent application is hereby incorporated herein by reference. 
     
    
     FIELD OF THE INVENTION 
       [0002]    This invention relates to medical apparatus and procedures in general, and more particularly to systems for providing radiation therapy to a patient and use of the same. 
       BACKGROUND OF THE INVENTION 
       [0003]    Radiation therapy has been used for many years to treat a patient by radiating selected anatomy so as to destroy tumors and the like. 
         [0004]    Starting in the 1930&#39;s and continuing into the 1970&#39;s, “orthovoltage” X-ray therapy units of approximately 300 kVp/1.2 mm Cu HVL were widely used to treat malignancies (e.g., tumors). The word “orthovoltage” was based on the concept of “the right energy”, in the sense that a 300 kVp electrical potential, resulting in an approximately 160 keV X-ray beam, generates a photon energy such that the skin absorption and the bone absorption becomes approximately equal, which is a necessity for radiotherapy treatments within the skull, and an advantage for radiotherapy treatments elsewhere in the body. Energies lower than approximately 160 keV will deposit energy preferentially in bone. In this respect it should be appreciated that Compton scatter of the radiation within the bone effectively concentrates the radiation dose in the vasculature of the bone, which can destroy the bone vasculature. Thus, energies below approximately 160 keV are generally detrimental to the bone. 
         [0005]    The 300 kVp units used into the 1970&#39;s had the disadvantage of excessive skin dose delivery (e.g., when compared to the radiation dose which is actually delivered to tumors in the body, after taking into account the absorption of the radiation by intervening tissue). This is particularly true when the radiotherapy is delivered using 1, 2 or 3 fixed entrance “portals”. In such a situation, the skin dose may be 3 to 5 times the tumor dose, and the skin is generally destroyed before the tumor is destroyed. Multiple fixed oversized entry portals, and multiple radiation sessions, were used as a partial solution to the problem of excessive skin dose delivery. However, it should be appreciated that the imaging of tumors at that time was comparatively crude, limited to perhaps 10 mm or so accuracy, and was severely hampered by the lack of precise knowledge of the tumor edges in 3D. 
         [0006]    As a result, practitioners generally resorted to stationary multiple entrance portals, multiple radiation sessions and generally oversized beams to deal with the uncertainty of tumor location. The locations of the entry portals, and their sizes, were generally determined by “best effort” projections based on 2D X-ray images. See Joseph Selman, MD, “Basic Physics of Radiation Therapy,” 1960, “Therapy Planning” pp 214-243. 
         [0007]    It was recognized that higher photon energies could be used to address the known fall-off in radiation intensity that occurs with depth, whereby to improve the efficacy of the radiation treatment and spare the skin from the aforementioned radiation overdose problem. 
         [0008]    To this end, efforts were made to create higher energy X-ray tubes. However, the engineering problems associated with creating direct X-ray tubes over 300-400 kVp were substantial. Several 1 MeV units were built, however, these high energy X-ray tubes generally used Van de Graaf generators (see, for example, the work of High Voltage Engineering and Massachusetts Institute of Technology) or RF resonant systems (see, for example, the Maxitron systems of General Electric), and were often 20 to 30 feet long. 
         [0009]    The 1 to 2 MeV X-ray systems based on Van de Graaf generators and/or RF resonance beams generally fell out of favor due to the difficulty of moving the X-ray beams in 3D (e.g., because of the large physical size of the X-ray source enclosure). 
         [0010]    In the mid-1950&#39;s, the “Gamma Knife” system (e.g., of Leksell) introduced the idea of using multiple isocentric beams, at higher energy levels, to provide radiotherapy to the patient. More particularly, the Gamma Knife system provided hundreds of high energy (approximate 1.2 MeV) spherically-arrayed cobalt beams focused to an isocenter. These beams were (and are) poorly defined due to the large size of the cobalt source (typically 10&#39;s of mm) and the short collimators which are inherent in the spherical array. This poor beam definition is problematic since sharp beam edges are necessary in order to avoid unintentionally radiating nearby healthy tissue. In addition, the Gamma Knife system provides stationary beams, and utilizes a fixed circular shape, thereby providing only spherical treatment volumes. Thus, the Gamma Knife system is unable to provide the variable, controllable treatment volumes necessary to closely match the shape of the anatomy which is to be treated. In addition, the Gamma Knife system suffers from diffuse scatter of the primary beam, which is inherent in energy levels over about 1.2 MeV. 
         [0011]    The 1.2 MeV cobalt systems, including the Gamma Knife systems, generally fell out of favor largely due to their relatively large cobalt radiation sources (typically tens of mm) which caused the edges of the radiation beam to be very poorly defined, even with good collimators. This is due to the simple geometric problems inherent in collimating a large (e.g., 20 mm) radiation source (see  FIG. 1 ) versus collimating a small (e.g., 1 mm) radiation source (see  FIG. 2 ). 
         [0012]    More particularly,  FIG. 1  shows a typically large cobalt source  5  of about 20 mm size, such as is commonly used in the Gamma Knife system, being collimated by a collimator  10 . As can be seen from  FIG. 1 , it is not possible to collimate such a large source to a sharp edge (see “large error”  15 ), thereby effectively eliminating any possibility of providing a sharp clinical edge to the radiation beam.  FIG. 2 , on the other hand, is a schematic view showing the sharp edge resulting from a 1 mm X-ray source (or focal spot)  20  as it is collimated by a collimator  25 , and the resulting “small error”  30 . Note that with the large cobalt source  5 , the penumbra of the beam edge may take 10 mm or more to “fade out” radially, from the nominal or desired edge, thereby making the beam emitted by a large cobalt source effectively useless if the desired treatment beam is to be 10 mm or less. This is a significant problem, given the current interest in small, accurate treatment volumes. 
         [0013]    Subsequently, there was developed the linear particle accelerator (Linac) systems which are common today. These Linac systems typically have an energy level of about 6 MeV, which makes them highly effective in addressing the known fall-off in radiation intensity with depth, whereby to improve the efficacy of the radiation treatment and sparing the skin from radiation overdose. Although the secondary scatter envelope edge of a 6 MeV beam may be 4 to 5 times that of the early 300 kV beams (due primarily to in vitro tissue scatter, and not beam geometry), the secondary scatter problem was largely lost in the early days of Linacs due to the relatively low accuracy of tumor location which was common at the time. Note that in vitro tissue scatter of the radiation beam (which can reach 8-10 mm beyond the edge of a high energy radiation beam) cannot be suppressed by improved collimators. 
         [0014]    Concurrent with the aforementioned improvements in radiotherapy systems, anatomical imaging systems have also been significantly improving, such that a 1 mm location accuracy is now routinely obtained. As a result, the X-ray beams can become smaller, and they can be more precisely positioned with the use of improved imaging modalities (e.g., CT, MRI and PET). 
         [0015]    As a result, the tissue scattering which is inherent in high energy beams such as the current  6  MeV Linac systems, and which was previously of less concern due to the relatively low accuracy of tumor location, has now become a considerable problem. 
         [0016]    More particularly, the secondary scatter effects of a high energy beam causes a “scatter” or “fog” of radiation that may have a range of up to 8-10 mm outside of the actual geometric beam. See  FIG. 3 , which shows the significant radiation scatter which can occur within tissue when using a high energy (e.g., 6 MeV) beam, and  FIG. 4 , which shows the nominal beam scatter which can occur within tissue when using a low energy (e.g., 100 kVp-800 kVp) beam. More particularly,  FIG. 3  is a schematic view showing a typical 6 MeV Linac beam generated from an X-ray source  35 , with secondary radiation  40  (generated by scatter interactions within the body) extending out to the dotted line  45 . This is approximately 10 mm away from the geometric beam, generating the “scatter error”  50 . Thus, even with excellent beam collimation, a 6 MeV beam is significantly scattered in tissue, again eliminating the achievement of a sharp edge. By comparison,  FIG. 4  is a schematic view showing a 1.0 mm X-ray source  55 , operating at 100-800 kVp so as to generate only short-range “Compton scatter”  60  and hence generating a sub-1 mm scatter error  65 . 
         [0017]    As noted above, the secondary scatter effects of a high energy (e.g., 6 MeV) beam did not present a significant problem with early Linac systems, inasmuch as tumor location was considerably less accurate than it is today, and inasmuch as the beams were fairly large and the gains in reducing excessive skin dose significantly outweighed “actual beam edge” definition drawbacks. It was accepted as the price of a large gain in efficacy and a significant reduction in skin damage. However, it is obviously a severe limit for the small, accurate beams desired in radiotherapy today. Others aware of these problems have recognized the medical physics at work here, and have outlined the potential benefits of “older” orthovoltage beams (e.g., 300 kVp beams) for higher net accuracy. See “Intermediate energy photons (1 MV) to improve dose gradient, conformality, and homogeneity: Potential benefits for small field intracranial radiosurgery”, Brian M. Keller, et al, Med. Phys. 36 (1), January 2009, © 2009 Am. Assoc. Phys. Med.; “Electron and photon spread contributions to the radiological penumbra for small monoenergetic x-ray beam (≦2 MeV), 
         [0018]    Jean-Philippe Pignol, et al, Journal of Applied Physics 105, 102011 (2009), © 2009 American Institute of Physics; and “Experimental measurement of radiological penumbra associated with intermediate energy x-rays (1 MV) and small radiosurgery field sizes”, Brian M. Keller, et al, Med Phys 34 (10), October 2007, © 2007 Am. Assoc. Phys. Med 3996. 
         [0019]    A second major drawback associated with Linac systems, in particular for neuroradiosurgery, stems from the Linac&#39;s intended and historic advantage of high energy (and hence low fall-off as the beam travels within tissue). The high energy (e.g., 6 MeV) of the treatment beam causes the beam to continue on through the tissue at close to full destructive energy on the far side of the treatment zone, and the beam is increasing in diameter, and the “fog” of radiation still surrounds it, potentially wreaking havoc on normal tissue beyond or beside the treatment volume. This, combined with diffuse edges, pair production, and several other secondary effects associated with high energy beams, inherently limits what can be done to obtain a sharp beam edge at 6 MeV. By way of example but not limitation, with current Linac systems, even using multi-leaf collimators (2 to 3 mm are the smallest size that can work effectively) to tailor the 6 MeV beam, only 5 mm to 10 mm of beam precision is possible inside the body. This is ten times worse control than imaging accuracy would otherwise allow, and can prevent treatment of small lesions near critical anatomy such as the optic nerve. 
         [0020]    Much of the recent prior art and current developments are geared to somehow gaining control of the Linac “fog of radiation” problems. See U.S. Pat. No. 8,280,001, issued Oct. 2, 2012 to Wang et al. From this observation, it is pointless to make a multi-leaf, “fine grain” collimation system, capable of a 1 mm geometric beam formation, if scatter within the patient will still be 5 to 8 times greater than the 1 mm geometric beam. 
         [0021]    The presented X-ray edge definition problem is somewhat analogous to the situation where a very fine line (e.g., less than 1 mm) is to be drawn on paper—the pen used to draw it must be sharp (i.e., also less than 1 mm). 
       SUMMARY OF THE INVENTION 
       [0022]    It has occurred to the Applicant that the Orthovoltage (or mid-voltage) X-ray intensity provides this fine edge, inasmuch as essentially all tissue absorption is short range Compton scattering; but to use this mid-voltage X-ray intensity clinically, for very small, accurate beams, a new approach is needed. In other words, Applicant has determined that properly applied mid-voltage beams (e.g., in the range of 100 kVp to 800 kVp) still possess very desirable inherent edge accuracy properties for radiotherapy or radiosurgery, despite the widespread (and perhaps incorrect) opinion that mid-voltage beams are obsolete, and that “Linac is better”. However, Applicant also recognizes that a new approach is needed to use these “mid-voltage” X-ray beams so as to obtain their advantages without encountering their disadvantages. 
         [0023]    In particular, the present invention employs the concept of “critical limits” in defining how to use the considerable advantage of the all-Compton absorption of mid-voltage X-ray beams, leading to a closely-defined overall system for correctly applying mid-voltage beams to clinical problems. 
         [0024]    The advantages of mid-voltage beams to be utilized, and their useful ranges in the new system approach follow, presented as critical limits to a clinically-useful mid-voltage system: 
         [0025]    Critical Limit  1 . For beams between 100 kVp and 800 kVp (i.e., a “mid-voltage” beam), the absorption process in tissue is almost entirely Compton absorption. This is well understood, and a nominal scatter range can be assigned to it, generally from sub-1 mm to 1 mm, dependent upon the energy of the beam, which then delivers a known accuracy to the beam edge. By comparison, the 6 MeV Linac system is characterized by an edge scatter or “fog” which is approximately 5 mm to 8 mm wide, due to the many non-Compton processes taking place in the tissue at 6 MeV. These non-Compton processes generate long range scatter, which is essentially harmful radiation, which extends well outside the desired beam width. 
         [0026]    Therefore, the present invention preferably operates between 100 kVp and 800 kVp. 
         [0027]    Critical Limit  2 . Compton processes are up to 1.2 times as lethal to cells per Gy as non-Compton processes, although that also applies to any errant radiation. See Joseph Selman, MD, “Basic Physics of Radiation Therapy,” 1960, pp 430-438. Therefore, close definition of the mid-voltage beam is essential, and this is achievable using a radiation source of small size, between approximately 0.6 mm and 3.0 mm diameter, and preferably between approximately 0.6 mm and 1.5 mm, and preferably approximately 1.0 mm diameter. 
         [0028]    Critical Limit  3 . It is possible to design an X-ray tube in the mid-voltage energy range, providing the intended small focal spot or source size (e.g., of only 1.0 mm or so), by using a shallow anode angle (e.g., less than 16 degrees, and preferably between 8 and 16 degrees). 
         [0029]    By combining the foregoing Critical Limits  1 ,  2  and  3  (i.e., mid-voltage beam intensity with all-Compton absorption, close beam definition using a small source size, and beam generation via a shallow anode angle X-ray tube), with proper system geometry (e.g., a 660 mm source-to-tumor distance, and a multi-leaf collimator (MLC) made of 1.0 mm to 2.0 mm leafs or fingers, located approximately two-thirds of the 660 mm toward the patient), a beam edge definition of about sub-1 mm to about 1 mm results, and is maintained inside the body. In general, it is desirable for the distance between the X-ray tube and the treatment volume to range between 400 mm and 750 mm. 
         [0030]    Working with the foregoing, the problem still remains of the excessive skin dose delivery associated with mid-voltage X-ray beams, which is due to the more rapid beam intensity fall-off in tissue due to the higher absorption rate of a mid-voltage beam. In order to be clinically effective, the skin dose delivery must also be maintained within critical limits. Resolution of this problem is essential to the success of a medical radiotherapy application using sub-MeV energy. The present invention addresses this problem as follows. 
         [0031]    In the case of neuro radiosurgery, a greatly simplified but “worst case” numerical example will be used to explain the unifying underlying operating concept of this new system that addresses skin dose and intensity fall-off within tissue. The numerical descriptions which follow are only approximate due to many technical variables, but they are intended to be close to clinical values. 
         [0032]    For small radiosurgical beams, one can envision a 15 mm beam width (i.e., about the size of a dime) to treat a 14 mm spherical tumor located “deep” within the brain. If that beam is stationary, and if the tumor is located 10 cm into the head, then for a 2 mm copper hVL (approximately 350 kVp), the radiation dose at the tumor is about 28 percent of the radiation dose at the dime-sized skin entrance portal. To make a truly worst-case scenario, a factor of one-fifth will be used (i.e., the radiation dose at the tumor is assumed to be about 20% of the radiation dose at the skin entrance portal). See Joseph Selman, MD, “Basic Physics of Radiation Therapy,” 1960; also published depth/dose curves. 
         [0033]    A typical lethal dose to the tumor volume might need to be 20 Gy; but applying more than 2 Gy to the skin has severe consequences. From the 5:1 drop-off ratio (i.e., skin dose:tumor dose) discussed above, a “worst case” total of 100 Gy must be “passed” through the skin portal in order to achieve the 20 Gy treatment dose on the deep-set tumor volume. If this is to be limited to 2 Gy at any one skin entrance portal (i.e., the desired maximum skin dose), there must be at least 50 different dime-sized entrance portals, with each entrance portal limited to 2 Gy each. 
         [0034]    It is obvious that, in fact, 50 “dimes” will fit on the general hemisphere of a human skull, so requiring the use of 50 different skin entrance portals is plausible for the single session destruction of a tumor of the aforementioned size via radiosurgery. All beams remain sharply defined, since (i) a collimated, small diameter (e.g., 1 mm) radiation source is used, whereby to eliminate the edge definition issues associated with large radiation sources, and (ii) a mid-voltage (e.g., 100 kVp to 800 kVp) radiation source is used, whereby to provide minimal (i.e., Compton-only) scatter of the beam in the tissue. This approach improves beam accuracy and hence also improves clinical outcomes. In addition, and significantly, the exit dose (i.e., the dose imposed on tissue located on the far side of the tumor) is far lower with the mid-voltage beam than is the exit dose of Linac-based 6 MeV beams, due to the more rapid fall-off of mid-voltage beams in tissue, a major issue for head and neck treatments. 
         [0035]    Thus, a significant aspect of the present invention is that the X-ray source and its associated collimator produce a small diameter beam, such that 50 or so skin entrance portals may be used to direct radiation at the treatment volume. In general, it is preferred that the X-ray beam be 30 mm or less in diameter. In one preferred form of the invention, the X-ray beam has a diameter of 15 mm. In another preferred form of the invention, the X-ray beam has a diameter between 2 mm and 25 mm, corresponding to the projected small tumor size. 
         [0036]    It will be appreciated that the small diameter, well-defined mid-voltage X-ray beam may be directed to the treatment volume through the 50 or so skin entrance portals by delivering discrete X-ray dosages at each of the skin entrance portals or by scanning a continuously-emitting X-ray source across the range of skin entrance portals; of course, for higher keV or other geometries, less than 50 portals will suffice. 
         [0037]    More particularly, in one form of the invention, the X-ray beam may be directed to the treatment volume by positioning the X-ray source in line with one skin entrance portal and then activating the X-ray source so as to deliver therapeutic radiation to the treatment volume while limiting harmful radiation to the skin (e.g., delivering approximately 0.4 Gy to the treatment volume and no more than 2 Gy to the skin). Then the X-ray source is moved so as to be in line with another (e.g., adjacent) skin portal and radiation is delivered to the treatment volume. This process is continued until the appropriate treatment dose has been delivered to the treatment volume. 
         [0038]    Alternatively, and more preferably, in another form of the invention, the X-ray beam may be directed to the treatment volume by scanning a continuously-emitting X-ray source across the range of skin entrance portals so as to deliver therapeutic radiation to the treatment volume while limiting harmful radiation to the skin (e.g., delivering approximately 20 Gy to the treatment volume and no more than 2 Gy to any one skin entrance portal). 
         [0039]    The actual location of the 50 different entrance portals (i.e., the “50 dimes”) is set by a medical planning protocol, taking into account what tissue is to be radiated and what tissue is to be avoided. Preferably this protocol is effected through the use of computerized, 3D dose planning software, which will be apparent to those skilled in the art in view of the present disclosure. Significantly, this dose planning software can take advantage of the sharp edge definition, and rapid fall-off, of the mid-voltage beam. 
         [0040]    It will be appreciated that the exact number of different entrance portals utilized is a function of the location of the treatment volume, the intervening tissue, the strength of the X-ray beam, the width of the X-ray beam, etc. In general, the net area of the entrance track for the X-ray beam (i.e., the sum of all of the different skin entrance portals) is approximately 20 to 100 times the beam area at a given skin entrance portal. 
         [0041]    It should be appreciated that the mid-voltage beams may be circular in cross-section, or they may be non-circular in cross-section, or they may be dynamically shaped through the use of multi-leaf collimators (MLC), e.g., with 1-2 mm leafs, wherein the shaping of the beams is dictated by the tumor shape in any given projection (e.g., using CT data for any given projection). 
         [0042]    The advantage of a mid-voltage beam having closely-defined edges remains throughout the procedure, as the angle of projection of the beam, and its cross-sectional configuration (adjusted by varying the MLC apertures) change to radiate a given projection or view. 
         [0043]    There is, therefore, adjustability to the angle of projection of the closely-defined beam, and to the cross-sectional configuration of the closely-defined beam, so that the system can be used to effectively match the radiated volume to the tumor volume. 
         [0044]    The present invention may be implemented by mounting an X-ray tube, capable of generating a point source, mid-energy beam, or multiples of such X-ray tubes, on one or more suitable mechanical gantries, such that the X-ray tube (or X-ray tubes) can be adjustably directed toward the isocenter of the treatment volume, e.g., so that the X-ray tube (or X-ray tubes) can be schematically positioned on the inner surface of an imaginary sphere (or other curved geometry). By way of example but not limitation, a metallic or non-metallic C-arm may be provided, wherein the C-arm carries an X-ray source and an associated beam collimator that can move along the C-arm, while centered on the treatment volume, such that the X-ray source can direct its beam to the treatment volume using a plurality of different skin entry portals. The C-arm may be movably mounted on bearings so as to provide a wider range of X-ray tube positions, and hence a wider range of skin entry portals. 
         [0045]    Although highly variable with X-ray tube design, filtration and energy, approximately 2 to 4 Gy per minute per source may be delivered to the treatment volume where the source is set 660 mm from the treatment volume and the treatment volume is located 10 cm inside the body. This level of radiation output would indicate a 25 minute treatment time with one X-ray tube source and a 10 to 15 minute treatment time with two X-ray tube sources, for delivering a 20 Gy treatment dose 10 cm inside the body. Of course, in accordance with the present invention, the 20 Gy treatment dose is delivered by using a plurality of skin entrance portals so that each skin entrance portal receives no more than 2 Gy at any one skin entrance portal. Two X-ray tube sources are generally preferred over more than two X-ray tube sources for a number of reasons, including less collision interference (between the X-ray tube sources) with only two trajectories. In accordance with the present invention, the distance between the X-ray tube source(s) and the target tissue is reduced to 660 mm, as compared to the traditional distance of 1.0 meter or 0.8 meter, in order to approximately double the amount of radiation delivered to the treatment volume per source, while still maintaining the beam resolution desired. The desired X-ray beam may be created by using a small focal spot source (e.g., about 1.0 mm) tailored with a collimator, e.g., a 1 mm MLC. The desired small focal point X-ray beam may be achieved by using a more acute X-ray anode angle, usable with intended small fields, and thus a smaller focal spot projection, while keeping power levels high (4-5 kW). Certain recent very high power rotating anode X-ray tube designs now used in CT can be adapted to this concept by engineering effort; at present 25 kW steady state is possible which can reduce these example times by a factor of 5× or 10×. 
         [0046]    The patient table, and thus the tumor location (which is indexed to the table by known stereotactic methods, and possibly checked by imaging components on the same X-ray system), is adjusted in an X-Y-Z, micrometer motion relative to the mechanical gantry, and hence relative to the X-ray source(s) movably mounted to the mechanical gantry. 
         [0047]    The present invention is intended to enable radiosurgical ablations with beams of 30 mm or less (and preferably of about 15 mm) in a manner inherently much more precise than Linac or Gamma Knife designs, and at a lower cost. Radiation shielding for 350 kVp beams is minimal compared to MeV devices, greatly lowering installed cost, while providing a better, more well-defined and more controllable bolus of radiation. From a cost-of-treatment perspective, 3 or 4 systems of the present invention can be installed for the cost of a single Linac system. 
         [0048]    At the same time, in addition to providing radiotherapy, the mid-voltage, small area X-ray sources (and collimators) can be used to provide exquisite X-ray imaging of the patient anatomy. Note that Linac and Gamma Knife designs do not provide this opportunity. By way of example but not limitation, a flat plate X-ray imager (e.g., such as the Paxscan flat plate X-ray imager made by Varian) may be placed on the far side of the anatomy, such that the X-ray source and the X-ray imager can be used to image the anatomy. 
         [0049]    In one preferred form of the invention, there is provided a radiotherapy system comprising an X-ray tube operating at 100 to 800 kVp for providing X-ray beams of 30 mm diameter or less and configured to move the entrance beam footprint on the body during irradiation to any arbitrary sequential position set that has been predetermined to limit the intervening tissue dose rate at any one location to a safe level, such that the sum of the skin area traversed during treatment is 20 to 100 times the beam area. 
         [0050]    In another preferred form of the invention, there is provided a radiotherapy system, the system comprising: 
         [0051]    a gantry; 
         [0052]    at least one X-ray tube movably mounted to the gantry, wherein the at least one X-ray tube operates at 100 kVp to 800 kVp and has a beam origin of 1.5 mm or less; 
         [0053]    a collimator associated with the at least one X-ray tube for collimating the X-ray beam from the X-ray tube to 30 mm or less; 
         [0054]    wherein the gantry and the at least one X-ray tube are configured so that (i) the X-ray beam from the at least one X-ray tube is always centered on a treatment volume, and (ii) the skin entrance portal of the X-ray beam may be moved on the body during therapy so that the total area of the skin entrance portals is 20 to 100 times the X-ray beam area. 
         [0055]    In another preferred form of the invention, there is provided a method for providing radiotherapy to a treatment volume, the method comprising: 
         [0056]    providing a radiotherapy system, the system comprising:
       a gantry;   at least one X-ray tube movably mounted to the gantry, wherein the at least one X-ray tube operates at 100 kVp to 800 kVp and has a beam origin of 1.5 mm or less; and   a collimator associated with the at least one X-ray tube for collimating the X-ray beam from the X-ray tube to 30 mm or less;   wherein the gantry and the at least one X-ray tube are configured so that the X-ray beam from the at least one X-ray tube is always centered on a treatment volume; and       
 
         [0061]    radiating the treatment volume using the X-ray beam, with the X-ray tube being moved so that the skin entrance portal of the X-ray beam is moved on the body during therapy so that the total area of the skin entrance portals is 20 to 100 times the X-ray beam area. 
     
    
     
       BRIEF DESCRIPTION OF THE DRAWINGS 
         [0062]    These and other objects and features of the present invention will be more fully disclosed or rendered obvious by the following detailed description of the preferred embodiments of the invention, which is to be considered together with the accompanying drawings wherein like numbers refer to like parts, and further wherein: 
           [0063]      FIG. 1  is a schematic view showing a typically large cobalt source of about 20 mm size, as is commonly used in Gamma Knife systems—as can be seen from the figure, it is not possible to collimate such a large source to a sharp beam edge; 
           [0064]      FIG. 2  is a schematic view showing the sharp beam edge which can be provided by a 1 mm X-ray source (or focal spot); 
           [0065]      FIG. 3  is a schematic view showing a typical 6 MeV Linac beam, with secondary radiation generated by scatter interactions within the body extending out approximately 10 mm beyond the normal boundary of the geometric beam—thus, even if the 6 MeV Linac beam has excellent collimation, beam scattering within tissue eliminates any possibility of a sharp beam edge; 
           [0066]      FIG. 4  is a schematic view showing how a 1.0 mm X-ray source, operating at 100-800 kVp, generates only short-range Compton scatter, and hence generates a sub-1 mm to about 1 mm error beyond the geometric beam; 
           [0067]      FIG. 5  is a schematic view showing a radiotherapy system which comprises one preferred form of the present invention, wherein the radiotherapy system comprises a plurality of C-arms mounted to a vertical axis bearing, wherein each of the plurality of C-arms carries an X-ray source (and associated collimator) that can move along the arc of the C-arm while centered to a treatment volume; 
           [0068]      FIGS. 6-8  are schematic views showing construction details of an X-ray source and associated collimator of the sort which may be utilized in the radiotherapy system shown in  FIG. 5 ; 
           [0069]      FIG. 9  is a schematic view showing another form of radiotherapy system formed in accordance with the present invention, wherein the radiotherapy system shown in  FIG. 9  is generally similar to the radiotherapy system shown in  FIG. 5  except that the C-arms are mounted to a horizontal axis bearing; 
           [0070]      FIGS. 10 and 11  are schematic views showing another form of radiotherapy system formed in accordance with the present invention, wherein the radiotherapy system comprises an imaging device which allows imaging with the same X-ray tube used for providing radiotherapy; 
           [0071]      FIG. 12  is a schematic view showing another form of radiotherapy system formed in accordance with the present invention, wherein the radiotherapy system comprises an imaging device which allows imaging with the same X-ray tube used for providing radiotherapy, and further wherein the radiotherapy system comprises a gantry in the form of a rotatable parallelogram linkage that allows variable positioning of the X-ray source and imaging device relative to the isocenter; 
           [0072]      FIGS. 13 and 14  are schematic views showing another form of radiotherapy system formed in accordance with the present invention, wherein the radiotherapy system is provided in conjunction with a CT machine, and further wherein the radiotherapy system comprises a single C-arm movably carrying two X-ray sources thereon; and 
           [0073]      FIGS. 15-18  are schematic views showing another form of radiotherapy system formed in accordance with the present invention, wherein the apparatus is generally similar to the apparatus shown in  FIGS. 13 and 14  except that the radiotherapy system comprises a two C-arms each movably carrying one X-ray source thereon. 
       
    
    
     DETAILED DESCRIPTION OF THE PREFERRED EMBODIMENT 
       [0074]    Looking now at  FIG. 5 , in accordance with the present invention, there is provided a radiotherapy system  70  which comprises two X-ray tubes  75  which are mounted to two C-arms  80  which are pivoted through generally vertical axis bearings with common centers  85 , aligned with the spherical isocenter  90 . X-ray tubes  75  are mid-voltage devices, operating at 100 kVp to 800 kVp, and have a small point source (e.g., 0.6 mm-3 mm). The two X-ray tubes  75  will always point, during a treatment, at isocenter  90 . The X-ray beams of X-ray tube  75  are collimated with collimators  95  so as to produce a small diameter (e.g., 30 mm or less, and preferably of about 15 mm), well-defined (e.g., with little collimation error) beam, such that 50 or so skin entrance portals may be used to direct well-defined beams of radiation at the treatment volume. Preferably collimators  95  are multi-leaf collimators (MLC), which need not operate symmetrically about the beam axis, such that collimators  95  can shift the net beam irradiated field to one side or the other of the X-ray tube central ray line. 
         [0075]    The rotation about the vertical axis  85 - 85  of C-arms  80  are driven by independent motors associated with the bearings  85  to precisely position the C-arms  80  rotationally around axis or center line  85 - 85 . 
         [0076]    X-ray tubes  80  and collimators  95  can “climb” up and down C-arms  80  under precise control of motors, using computer feedback methods. 
         [0077]    Multi-leaf collimators  95  may open and close as they travel, to track a known projection or shape of the tumor as viewed from their instantaneous perspective. This data is derived from 3D CT images in known ways, and this process is allowed for by design, by making multi-leaf collimators  95  fast enough to follow this. 
         [0078]    The axis  85 - 85  may be tilted, or the patient head tilted, to best access the treatment zone. 
         [0079]    The volume to be radiated is positioned to  90  by precise X-Y-Z motions of the table  100 . Table  100  may or may not move during treatment, under control of treatment planning software. X-ray tubes  75  are generally operated in this device at 100 kVp to 800 kVp, and preferably at 300 kVp to 500 kVp. They have a shallow anode angle of approximately 8 to 16 degrees to address high loading of a nominal 1 mm source size. They may be rotating anode tubes. 
         [0080]    Multi-leaf collimators  95  are preferably composed of about thirty-two 1.5 mm leafs on each of the four sides. Each leaf is about 15 mm tall, of tungsten, and each has its own drive nut screw, motor and feedback of position, so 128 motors are used per collimator. More or fewer motors can be used depending on the maximum beam cross-section desired, and the leaf dimension. By way of example but not limitation, and looking now at  FIGS. 6 and 7 , there is shown an exemplary collimator  95  mounted to an X-ray tube  75 . In one preferred form of the invention, collimator  95  comprises a pair of collimator assemblies  105 , wherein each of the collimator assemblies  105  comprises a plurality of leafs  110  which may be adjustably positioned within the opening  115  of a housing  120 . Preferably collimator assemblies  105  are set in series, and perpendicular to one another. 
         [0081]    X-ray tubes  75  are operated at about 4.5 to 5.0 kW on constant potential DC power supplies, and if of special known design, up to 5 times that. They are oil-cooled by circulating pumps. They have filters specific to the HVL Cμ desired, generally between 1.2 and 3.0 mm HVL, and may be multi-layer filters. 
         [0082]    For use in neuroradiosurgery, precision table  100  enters the arcs of C-arms  80  horizontally to position the head and tumor to within 1 mm of isocenter  90 . Various head frames and frameless approaches can be used to immobilize the head and register it into the overall system coordinate system, which should be accurate to about 1 mm. 
         [0083]    Thus it will be seen that, in accordance with the present invention, there is provided a radiotherapy system  70  which generally comprises a gantry (which generally comprises at least one C-arm  80 ) upon which at least one X-ray source (which preferably comprises at least one X-ray tube  75  and associated collimator  95 ) is movably mounted. The geometry of the gantry and the X-ray source are such that the X-ray source is always focused at an isocenter  90  regardless of the position of the X-ray source along the gantry; however, inasmuch as the X-ray source is movable along the gantry, the system allows the radiation beam emanating from the X-ray source to be directed at isocenter  90  from a variety of angles, whereby to provide a large number of different skin entrance portals. Preferably the gantry is movable relative to the isocenter (e.g., each C-arm  80  is pivotable about an axis) so as to increase the range of positions from which the X-ray source may be directed at isocenter  90 , and hence further increases the range of different skin entrance portals. In accordance with the present invention, each of the X-ray sources comprises a mid-voltage X-ray source (e.g., 100 kVp to 800 kVp) having a small beam origin (e.g., about 1 mm), such that when it is appropriately collimated by collimator  95  into a small diameter beam (e.g., 30 mm or less, and preferably of about 15 mm), provides a well-defined beam which is subject to minimal scattering within the tissue, whereby to provide a tightly constrained X-ray beam to the anatomy. In accordance with the present invention, there is also provided a motorized table  100  upon which the patient may be positioned, such that by precise movement of motorized table  100  relative to the gantry, the target tissue may be positioned at isocenter  90 . As a result of the foregoing construction, the narrow, well-defined, mid-voltage X-ray beams may be directed at the isocenter from a substantial number of different positions (e.g., 50), so as to deliver a therapeutic radiation dose to the isocenter while maintaining the radiation dose at each skin entrance portal below a harmful level. 
         [0084]    In  FIG. 5 , C-arms  80  are shown as pivoting through generally vertical axis bearings with common centers  85 . However, as seen in  FIG. 9 , C-arms  80  may alternatively be configured so as to pivot through generally horizontal axis bearings with common centers  85 . This configuration allows placement of X-ray tubes  75  along the long axis of the patient. 
         [0085]    Obviously, many other mechanical arrangements such as robotic arms, hemispherical tracks, spirals in a plane, etc., can be used to position the 100 to 800 kVp tubes relative to the patient. The multi-arc system shown in  FIGS. 5-9  show just some of the possible embodiments of the present invention. 
         [0086]    Looking next at  FIGS. 10 and 11 , imaging devices  125  (e.g., X-ray sensitive flat panels such as the Varian Paxscan 4030 of Varian Medical Systems, Salt Lake City, Utah) can be placed on the exit side of the beam trajectory by other known mechanical means, whereby to enable fluoro or cone beam CT, or projection imaging, even during treatment. The imaging plate may incorporate a movable protective cover or shield so as to protect the imaging plate during long treatments, and/or the energy may be reduced during imaging. 
         [0087]      FIG. 12  shows a parallelogram linkage  130  which is rotatably mounted to a support stand  135  and acts as the gantry for movably supporting X-ray tube  75  (and its associated collimator  95 ) and imaging device  125  relative to isocenter  90 . If desired, parallelogram linkage  130  may be of the sort described and illustrated in U.S. Pat. No. 3,892,967, issued Jul. 1, 1975 to John K. Grady et al. for APPARATUS FOR RADIOLOGICAL EXAMINATION OF A SUBJECT THROUGH A SOLID ANGLE, which patent is hereby incorporated herein by reference. 
         [0088]    Looking now at  FIGS. 13 and 14 , a CT scanner  140  can be placed in front of radiotherapy system  70  (comprising a single C-arm  80  movably supporting two 
         [0089]    X-ray tubes  75  thereon), and can be indexed to isocenter  90 , such that a calibrated CT scan can be performed, and then the patient shifted a precise amount, for treatment within the coordinate system. 
         [0090]    See also  FIGS. 15-18 , where CT scanner  140  is placed in front of a radiotherapy system  70  comprising two C-arms  80 , wherein each C-arm  80  movably supports one X-ray tube  75  thereon. 
       MODIFICATIONS OF THE PREFERRED EMBODIMENTS 
       [0091]    It should be understood that many additional changes in the details, materials, steps and arrangements of parts, which have been herein described and illustrated in order to explain the nature of the present invention, may be made by those skilled in the art while still remaining within the principles and scope of the invention.