The research for this invention was sponsored by the U. S. Army, grant number BC 99087. The U. S. government has certain rights to this invention.

Background of Invention Field of Invention This invention relates to a device for radiation treatment for cancer and the method thereof. It more specifically relates to a device to generate helical electron beams for radiation therapy for tumors and the method of irradiating tumors with a helical electron beam.

Description of the Related Art Radiation therapy has been effective for managing early stage (stage I-II) breast cancers following breast conservation surgery (BCS). More recent studies have indicated that radiation therapy also improves survival rate for late stage (stage III-IV) breast cancer patients. However, the high risk of cardiac or pulmonary complications often compromises the treatment of these patients using conventional radiation therapy techniques.

Conventional photon beams have been used successfully for treating early stage breast cancer treatment. It is now widely accepted that radiation therapy following breast conservation therapy is as effective as mastectomy for early breast cancers [1-9]. Radiation therapy has also been used for managing late-stage breast

cancer patients [10-16]. More recent studies have found that radiation therapy improves survival for a subgroup of patients with late-stage breast cancer [15-16].

Conventional radiation therapy techniques use two opposing tangential photon beams to treat the whole breast. The advantage of photon beam is that it has sharp beam delineation and high penetrating power. The disadvantages of photon beam treatment are as follows: (1) Limited beam setup options. Patient geometry may sometimes include excessive heart or lung volume for irradiation.

(2) Unable to treat internal mammary nodes because the proximity of the node chains to the heart.

(3) Potential high scatter dose to the contralateral breast.

(4) Patient breathing motion may introduce accidental irradiation of large volume of heart or lung.

These restrictions especially the first one have restricted the use of photon beam radiation therapy for advanced breast cancer treatment. Several clinical studies have reported significant rate of cardiac or pulmonary toxicity associated the conventional photon beam techniques [17-27]. High risk of irradiating the heart or lung becomes critical for post-mastectomy patients when they receive chest wall or internal mammary node treatments [20-21]. The limitation of photon beam breast treatment is evident from a recent study that compares the photon Intensity Modulation Radiotherapy (IMRT) technique with the conventional radiation therapy techniques for breast cancer treatment [28]. The results of the study show that IMRT presents almost no advantage over the conventional techniques for whole breast

treatment. This result implies that there may be not much room in improving the existing techniques for photon beam treatment of breast cancers.

Electron beams, on the other hand, offer attractive features for breast cancer treatment. First, breast tumors are relatively shallow that could be penetrated by electrons of 18-20 MeV. These electron beams are already available for most medical accelerators. Second, the electron beam exhibits fast fall-off near the end of its range of travel. These features enable treatment of the internal mammary nodes or the chest wall region while sparing the heart or lung underneath. Because the anterior-posterior direction of the breathing motion coincides with the beam-on direction, the simple treatment setup reduces the possibility of over irradiation of the heart or lung due to patient breathing motion. This is analogous to the treatment of lung cancer patients where direct anterior-posterior setup has been found to be the most reliable and effective treatment method [29]. In comparison, electron beam would overcome all the above listed disadvantages associated with the photon beam treatment. One major disadvantage of electron beams is their excessive scatter and beam penumbra.

Recent rapid developments in computer controlled beam delivery systems have become a driving force in developing intensity modulation radiotherapy (IMRT) using high-energy photon beam [30-49]. The technology has become mature enough that several clinical trials are near their completion [48-49].

Because of the high cost and few locations that can handle high-energy photon beam radiation therapy, it would be useful to have a different approach to reduce radiation to non-cancerous tissue while increasing the radiation to cancerous tissue.

One approach would be to use electron beams. But the problem with electron beams is their tendency to scatter.

If one could reduce the problems associated with electron beam radiation, one would be able to develop intensity modulated electron beam for Avoidance Radiation Therapy (ART). The idea of ART has been applied to almost every aspect of radiation therapy where a tumor dose is typically limited by the tolerance of normal tissue structures. If normal tissue dose could be significantly reduced or avoided, then the complication rate would be reduced. This reduction in radiation of normal tissue would permit dose escalation to achieve better local control of the tumor and improves the patient survival [50]. Several promising beam modulation techniques have been proposed for ART treatment using electron beams [51-57].

One approach is using finely collimated beams scanning across the entire treatment area using more sophisticated beam transport systems such as those used in the Racetrack Microtron [51-54]. This approach requires significant modification of the design for most clinical linear accelerators; and thus is not feasible. Another approach is to simply use the photon beam multileaf collimator system (MLC) for electron-beam intensity modulation [57]. This approach has not yet solved the problem the excessive electron beam scatter effects.

There have been several theoretical studies and preliminary experiments exploring the use of static magnetic fields in shaping clinical electron beams [58-63].

None of these studies and experiments has proven to be successful yet.

Several unique beam delivery algorithms using computer controlled MLC systems have been recently developed for beam delivery optimization and verification [39-47]. Conventional three-dimensional pencil beam algorithm is a fast and practical method used in most clinical treatment planning systems [64-65]. More time- consuming Monte Carlo method has also been recently developed particularly for

various heterogeneous cases [66-68]. These algorithms could be used or modified for an improved electron beam radiation system.

Brief Summary of the Invention It is an object of this invention to have a device for achieving intensity, energy and spatially modulated Avoidance Radiation Therapy for treatment of cancer. It is a further object of this invention that this device could be directly attached to the gantry of the medical accelerator similar to conventional electron cone or can be attached to a moveable platform or stand and detachably secured into position during radiation therapy. It is a further object of this invention that the device use a radiation source inside a linear accelerator to generate electrons and one or more magnets with the magnetic field in the same direction as the path of the electrons to cause the electrons to travel in a helical trajectory parallel to the magnetic field.

It is object of this invention to have a tertiary collimating cone with a dynamic energy compensator and a magnetic electron collimator which, when used in conjunction with a linear accelerator, induces electrons from the linear accelerator to travel in a helical pathway. It is a further object of this invention that the magnetic field of the magnetic electron collimator is parallel to the direction of the travel of the electrons. It is a further object that this device can be used to irradiate cancers or tumors without irradiating non-cancerous tissue or reducing the amount of radiation hitting non-cancerous tissue.

It is an object of this invention to inhibit electrons from scattering or diverging during radiation treatment of a patient. It is a further object to cause electrons to travel in a helical trajectory. It is another object to use magnets to induce the helical path of electrons.

It is an object of this invention to treat cancer patients, human or animal, with radiation in the form of electrons having a helical trajectory. It is an object of this invention to treat various cancers such as, but not limited to, breast cancer, melanoma, head and neck cancer, lymphomas, nasopharyngeal carcinoma, sarcomas in the extremities, and testicular cancer.

It is an object of this invention to conduct radiation therapy in patients having cancer wherein the radiation therapy uses a helical electron beam. It is a further object of this invention to treat cancer with a helical electron beam while avoiding unnecessary irradiation of the patient's non-cancerous tissue.

It is an object of this invention to have a device that emits a helical electron beam for conducting radiation therapy for cancer in patients. It is a further object that this device will assist in avoiding unnecessary irradiation of the patient's normal tissue.

It is another object of this invention to treat a patient with breast cancer and in need of treatment with a helical electron beam with a conformal dose distribution that is targeted to the whole breast, the chest wall, and/or the internal mammary nodes. It is a further object to prevent excessive irradiation of the heart and/or lung.

It is an object of this invention to have a device for generating helical electron beams with a conformal dose distribution for radiation therapy in order to treat whole breast, the chest wall, and/or the internal mammary nodes. It is a further object to this invention that this device will assist in preventing excessive irradiation of the heart and/or lung.

It is an object of this invention to have a device that when used in conjunction with a linear accelerator that the combination emits a helical electron beam. It is a

further object of this invention that the electron beam is collimated at or near the patient's skin surface. It is a further object of this invention that the electron beam is collimated between 0 cm to 40 cm from the patient's skin and most preferably between 0 cm and 10 cm from the patient's skin.

It is an object of this invention to have a device that when used in conjunction with a linear accelerator that the combination emits a helical electron beam. It is a further object of this invention that the device reduces the electron beam's divergence and penumbra. It is a further object of this invention that the shape of the field of the helical electron beam ranges between 0.01 cm2 to 800 cm2, and more preferably between 0.5 cm2 and 10 cm2.

It is an object of this invention to have a device for generating helical electron beams for radiation therapy whereby the field size of the treatment area is between 0.1 cm2 and 200 cm2 and more preferably between 0.2 cm2 and 1.0 cm2.

It is an object of this invention that the energy of the foil modulate the energy of the electrons passing through the foil. It is a further object of this invention that the modulated electrons have energy ranging between 0.5 MeV and 40 MeV and more preferably between 4 MeV and 20 MeV.

It is an object of this invention to have a device for generating a helical electron beam wherein the device uses a linear accelerator to generate the electron beam and a magnet with the magnetic field in a parallel path as the direction of travel of the electrons. It is a further object of this invention to use at least one foil to modulate the energy of the electrons.

It is an object of this invention to have a device for generating a helical electron beam wherein the device uses a linear accelerator to generate the electron

beam and a magnet with the magnetic field in a parallel path as the direction of travel of the electrons. It is a further object of this invention to use at least one foil to modulate the energy of the electrons. It is a further object of this invention to have a multileaf collimator adjacent to or within close proximity to the magnet. It is a further object of this invention to be able to move the magnet and the multileaf collimator quickly and easily into proper position for treatment.

Description of the Several Views of the Figures Figure 1 is a perspective view of the tertiary collimating system for generating helical electron beams.

Figure 2 is a cross sectional view an alternative embodiment of the helical beam generator.

Figure 3 illustrates the relative dose of the electrons at zero magnetic field, 20 kG magnetic field, and 50 kG magnetic field.

Figure 4 illustrates the dose enhancement of the electrons at various depths for zero magnetic field, 20 kG magnetic field, and 50 kG magnetic field.

Figure 5 illustrates the computed isodose dose distribution of an energy and intensity modulated electron beam for a chest wall treatment.

Detailed Description of the Invention Static magnetic fields can shape the electron paths. Under an axial magnetic field, any scatter or divergent electrons can be constrained to follow helical trajectories. The simulation of electron beam trajectories under an axial magnetic field indicates electrons traveling in a helical path. Electrons under an axial magnetic field have a much sharper penumbra and improved range straggling. However, strong magnetic fields about 3-6 Telsa are needed to generate a useful helical electron beam

with gyration radius of 1-2 mm [59]. Such strong magnetic fields require bulky and expensive super-conducting magnets.

As shown in Figure 1, the preferred embodiment of this invention uses a tertiary collimating cone, 50, having a magnetic electron collimator, 16, and a dynamic energy compensator, 3, to generate a modulated electron beam that travels in a helical path. The magnetic electron collimator contains a helical beam generator, 2, which uses a magnetic coil, 4, that produces a magnetic field direction that is aligned with the direction of movement of the electron beam. As a result, the solid angle of each individual modulated electron field is significantly reduced. This arrangement effectively decreases the axial magnetic field required for producing desired helical electron beams.

When electrons pass through a magnetic field, they are confined to travel in helical trajectories about an axis parallel to the magnetic field lines with a specific gyration radius. The gyration radius is determined by the following formula: a = pi (Formula 1)<BR> 3x10-7B where a is the gyration radius in cm, p is the momentum perpendicular to the magnetic field lines in MeV/c (c is the speed of light in vacuum), and B is the magnetic field strength in kilo-gauss (kG). Because of this confinement, the resulting fluence at the phantom or patient plane will be increased resulting in dose distributions that will potentially improve the treatments of various cancers such as,

but not limited to, breast cancer, melanoma, head and neck cancer, lymphomas, nasopharyngeal carcinoma, sarcomas in the extremities, and testicular cancer.

For example, if a patient is treated at 100 cm Source to Skin Distance (SSD) by superposing modulated electron fields up to 3 x 3 cm2, then only magnetic field of 3-5 kG instead of 3-5 Tesla is needed to produce helical electron beams with gyration radius of 1-2 mm. Such uniform magnetic fields are generated using regular Helmhotz magnetic coils, but any hollow-shape magnets can work. Because the needed field strength is dependent on the scanning field size and initial beam energy, magnetic fields can range from 0. 5 kG to 50 kG. For specific beast cancer or chest wall treatments, helical electron beams ranging from 10 MeV to 20 MeV use magnetic field modulations ranging from 1 kG to 100 kG. For superficial cancer such as melanoma, helical electron beams range from 0.5 MeV to 8 MeV use magnetic field strengths ranging from 0.5 kG to 100 kG. For head and neck cancers involving lymph nodes such as nasopharyngeal carcinoma, helical electron beams ranging from 5 MeV to 15 MeV use magnetic field strengths ranging from 10 kG to 200 kG. For other cancers such as soft-tissue sarcoma in extremities, it is preferable to use finely adjusted helical electron beams ranging from 4 MeV to 20 MeV with magnetic field strengths ranging from 0.5 kG to 200 kG.

As shown in Figure 1, in the preferred embodiment, this invention is a tertiary collimating cone, 50, which is mounted to the gantry, 51, of a typical linear accelerator at the gantry's attachment ports, 52. The linear accelerator contains various components that are well known in the art field. In summary, the linear accelerator contains an electron source, 53, which produces an electron beam, 54.

The electron beam passes through a scattering foil, 55, and may pass through a

primary collimator, 56, and a multileaf collimator, 12, prior to exiting the gantry. It is known in the art field that the position of and order of passage through the multileaf collimator and the primary collimator can vary within the gantry.

In an alternative embodiment, the tertiary collimating cone, 50, can be attached to a movable stand, instead of attaching the tertiary collimating cone to the gantry of the linear accelerator. In this manner, one can move the tertiary collimating cone in the correct position between the patient and the linear accelerator. Then one can detachably secure the stand and tertiary collimating cone to the floor, wall, ceiling, table, gantry, or other site, so that the tertiary collimating cone does not move during treatment.

Alternatively, various components of the tertiary collimating cone can be added to the inside of the gantry. In this alternative embodiment, the magnetic electron collimator, 16, and dynamic energy compensator, 3, can be placed inside the gantry. Alternatively, the helical beam generator, 2, can be placed inside the gantry in line with the path of the electrons that pass through the primary collimator and the multileaf collimator. Foils or other producer of modulated electron can be placed between the collimators and the helical beam generator.

As shown in Figure 1, in the preferred embodiment, the tertiary collimating cone, 50, contains a dynamic energy compensator, 3, and a magnetic electron collimator, 16. The dynamic energy compensator, 3, contains two foil changers, 11, foils, 15, and a stepping motor, 10. The foil changers are located across from each other at the entrance of the tertiary collimating cone, 15. The foils, 15, are held by and stored within the foil changers, 11. While it is preferable to have at least fifteen foils in the dynamic energy compensator, more or less foils can be used. It is even

possible to not have any foils, if the energy of the electrons is appropriate. The actual number of foils used can vary based on the energy of the electron beam and the desired final energy of the helical electron beam. The foils can range in thickness between 0.5 mm to 5 mm, and most preferably range in thickness between 1 mm to 2 mm. The foils can be made from aluminum, tungsten, tantalum, any medium atomic number metals, and/or alloys of any medium atomic number metals. The stepping motor, 10, moves each foil independently of each other. The foils block the electron beam, 54, dynamically during a treatment to produce modulated electrons, 52. One problem that may occur is that the electrons can strike the foil and create X-rays.

This potential production of X-rays can be reduced to an insignificant amount by using mesh type foils and/or having the foils made from different metals or alloys.

Foil thickness can also help alleviate the potential production of X-rays.

Modulated electrons have different properties than the electron beam, 54, from the electron source, 53. The electron beam loses about 100-800 keV after penetrating one sheet of foil. The dynamic energy compensator, 3, effectively produces modulated electron beams that can range between 0.5 MeV and 40 MeV, but most preferably varying from 4 MeV to 20 MeV.

The magnetic electron collimator, 2, is located downstream from the dynamic energy compensator, 3, in the tertiary collimating cone, 50. The magnetic electron collimator has a dynamically focused helical beam generator, 2. The helical beam generator is a magnetic coil, 4, which produces an axial magnetic field, relative to the path of the modulated electrons.

In the preferred embodiment, the magnetic coil, 4, contained within the helical beam generator, 2, produces axial magnetic field ranging from 0.5 kG to 200 kG,

depending on the type of cancer or tumor being treated. The inner diameter of the magnetic coil in the helical beam generator can vary. The preferred inner diameter opening is between 0.1 cm and 30 cm, and most preferably between 1 cm and 20 cm.

In the preferred embodiment, the tertiary collimating cone has a locator controller which controls the location of the helical beam generator. In the preferred embodiment, the helical beam generator is attached to a focusing rail, 5. A driving motor, 6, is attached to and moves the focusing rail. As the focusing rail moves, the helical beam generator also moves. The driving motor is located on the exterior of the tertiary collimating cone and uses a stepping motor to move the focusing rail. The driving motor moves the focusing rail, and thus the helical beam generator into the desired position for optimal focusing of the electrons on the area of treatment. The helical beam generator can be located at any position along the bottom surface of the tertiary collimating cone. The helical beam generator can rotate 180° ; the axis of rotation being perpendicular to the plane of the patient's skin in the area of treatment.

A computer, 7, controls the driving motor and thus the movement of the helical beam generator along the focusing rail. The computer contains software that uses predetermined algorithms to determine the optimum position of the helical beam generator, the optimum dosage, and the optimum treatment protocol. This arrangement permits quick repositioning of the helical beam generator during radiation therapy sessions, thereby permitting small fields of treatment and treatment of many fields in one radiation therapy session. The radiation field size can range between 0.1 cm2 and 200 cm2, and most preferably between 0.2 cm2 and 1.0 cm2.

The speed of the computer, the algorithm, and the precision of the stepping motor all play a role in determining the size of the radiation field.

In an alternative embodiment of this invention as illustrated in Figure 2, a multileaf collimator, 12, is present within the magnetic electron collimator, 16. While Figure 2 illustates a multileaf collimator upstream of the magnetic coil, 4, the multileaf collimator can also be located downstream, or both upstream and downstream of the magnetic coil. Any of the various well known in the art field types of multileaf collimators can be used. In this alternative embodiment, the linear accelerator can either have a multileaf collimator or lack a multileaf collimator.

Because electrons are being modulated rather than photons, the multileaf collimator within the magnetic electron collimator, 16, required a much smaller leaf thickness than is required for other types of particles. In this alternative embodiment, the leaves of the multileaf collimator can range in thickness between 0.3 cm to 5 cm, and most preferably between 0.5 cm and 3.0 cm. This reduction in thickness in the multileaf collimator within the magnetic electron collimator significantly reduces the overall weight of the tertiary collimating cone, 50. Placing a multileaf collimator within the tertiary collimating cone provides for better focusing of the electron beam on the tumor and enables one to prevent applying the helical electron beam onto normal tissue. In this alternative embodiment, the multileaf collimator within the magnetic collimating cone also plays a role in determining the size of the radiation field.

The helical electron beam exits the helical beam generator located along the bottom surface of the tertiary collimating cone prior to striking the patient. The distance from the bottom surface of the tertiary collimating cone to the patient's skin can range from 0 cm to +40 cm, and most preferably from 0 cm to 10 cm. It is possible that part of the patient can enter inside the helical beam generator in which

case that area of the patient is located a negative distance from the bottom surface of the tertiary collimating cone.

Because this invention collimates the electron beam at or near the patient's skin surface, the invention reduces the electron beam's divergence and penumbra.

A computer, 7, assists in determining the optimal position, moving the helical beam generator to the optimal position, and verifying that the helical beam generator is in the correct optimal position. The computer receives information from and transmits information to the helical beam generator via a data interface controller, 8, and a data bus, 9. The data interface controller and data bus also relay information from the computer to the stepping motor, 10, which controls the movement of the foil changer, 11, and back to the computer. The data interface controller and data bus also relay information from the computer to the multileaf collimator, 12, and back to the computer. The computer also assists in determining the optimal field size for the radiation treatment. If a multileaf collimator is located inside the tertiary collimating cone, 50, that multileaf collimator is also under the control of the computer, 7.

An algorithm, similar to photon IMRT beam delivery algorithms, is stored within the computer, 7. The algorithm assists in determining the correct position of the helical beam generator, 2, and the dynamic energy compensator, 3. The algorithms are based on the criteria of minimum total beam-delivery time and maximum dose accuracy. The algorithm uses a user-developed treatment planning system with graphical user interface.

The computer also contains an interface for quality assurance. For each patient, a graphical time-dose sequence based on patient CT data and the beam delivery sequence control data is produced to display the accumulated dose delivered

to the patient. This procedure allows accurate verification of the time-dose relationship. If any interruption happens during the treatment, the actual delivered dose is recovered from the actual beam-on time. Any leftover sequence is then compensated accurately once the interruption is cleared.

Software is used to simulate the time-dose sequence based on patient CT data and treatment plan generated from HEART delivery control files. For each patient, the time-dose sequence for the entire delivery is simulated using computer software for pretreatment quality assurance of a treatment delivery. Therefore, any interruption during the treatment delivery allows the operator to recover the interruption based on the simulated sequence to within 2% of the desired dose.

In the preferred embodiment of this invention, one plans radiation dose coverage of a given target volume by first segmenting the entire depth of the target into approximately fifteen discrete distances, each corresponding to a modulated beam energies. The shape of the target in electron beam's eye view for each distance is back-projected to determine the port of an intensity modulated electron beam for the select beam energy. The computer, using an algorithm based on the measured data table such as percentage depth dose (PDD), determines the beam intensity profile to cover the target at the desired depth. A standard least-square minimization method such as conjugate gradient search is used to optimize the weights of the intensity- modulated beams to achieve uniform and conformal dose coverage to the target.

While this embodiment for radiation dose coverage is preferred, other beam optimization algorithms which are well known in the art field, such as simulated annealing, linear and non-linear programming techniques can be used.

The overall size and shape of the tertiary collimating cone can vary. It may be preferable that its height range between 20 cm to 80 cm and that it measures between 20 cm to 80 cm from side to side.

For radiation cancer therapy, this invention permits greater accuracy and efficiency in doses, ranges of doses, depth of treatments, duration of treatments, and number of fractions. This invention also permits better targeting of radiation and highly reduces the amount of scattering of radiation outside the tumor tissue into normal tissue. This invention permits one to have more conformal treatment of tumor while sparing normal tissue from the radiation.

A Monte Carlo study was performed using the preferred embodiment of this invention. In this study, 15 MeV electrons were generated by an EGS4/BEAM simulation of an Elekta SL20 linear accelerator with the standard, known in the art field 10 x 10 cm2 electron cone (Elekta Oncology Systems Ltd., West Sussex, UK) attached to the end of the linear accelerator gantry. A version of the EGS4/DOSXYZ code was modified to generate the magnetic field module developed by Nelson and Rogers et al. [67-68]. In the simulation, a magnetic field with field lines parallel to the beam direction was applied in the region of air extending 10 cm immediately beyond the standard electron cone. (It is to be understood that in the invention, the magnetic field will be generated inside the tertiary collimating cone and can extend beyond the end of the tertiary collimating cone.) In the simulation, the magnetic field strengths used varied from 0 kG to 50 kG. A 100-cm3 water phantom for scoring dose was placed at the end of the 1 0-cm region.

With no magnetic field (0 kG) the electrons diverge to more than-7.5 cm and +7.5 cm from the center of the travel path as they travel further away from the end of

the electron cone resulting in a broader field at the phantom surface. With magnetic fields of 10 kG and 50 kG the electrons are forced to move in a"tunnel"with dimensions defined by the 10 x 10 cm2 aperture of the electron cone. In the 10 kG magnetic field, the electrons diverged approximately-6.25 cm and +6.25 cm from the center of the travel path. In the 50 kG magnetic field, the electrons barely diverged out of the tunnel of the 10 x 10 cm2 path defined by the electron cone. Using a 10 kG or 50 kG magnetic field, a more constrained dose distribution is achieved in the water phantom.

The isodose contours in the xz plane of the water phantom with no magnetic field (0 kG) and for a magnetic field of 10 kG and 50 kG indicate that with increasing magnetic field strength, the distribution of the electrons becomes more compact. The penumbra region decreased between approximately 10% to 50%, depending on the field strength. The higher the magnetic field, the smaller becomes the penumbra.

The penumbra is defined as between 20% to 80% of the dose profile.

Another effect of the magnetic field is the enhancement of the dose distribution as a function of depth within the water phantom. Figure 3 shows the central-axis depth dose curves for three cases (no field (-), B = 20 kG (----), and B = 50 kG ()). The curves are all normalized to the dose at the surface. In Figure 3, with no magnetic field (-), the dose distribution reaches a maximum of 1.15 at a depth of 2.5 cm. In a 20 kG magnetic field (----), the dose distribution is enhanced to a maximum of 1.25 at a depth of 4 cm. With a 50 kG magnetic field (), the dose distribution is enhanced further to a maximum of 1.32 at a depth of 4.2 cm.

The dose enhancement E is defined the following formula : (Formula2)<BR> D (d, B = 0) where D (d, B) is the dose at depth d with a magnetic field strength B applied to the 10- cm region in air just beyond the cone. Figure 4 shows E as a function of d for B = 20 kG (----) and B = 50 kG (), as compared to no magnetic field (-). With a magnetic field of 50 kG, the maximum dose enhancement occurs between 3-4 cm, declines slightly from 4 cm up to 7 cm and then approaches zero at approximately 7 cm. With a magnetic field of 20 kG, the maximum dose enhancement occurs between 4.5 cm and 5.7 cm and is equal in dose enhancement to the dose enhancement in the 50 kG magnetic field. The dose enhancement drops and regains strength in the 20 kG magnetic field between 5.8 cm and 6.5 cm then approaches zero at approximately 7 cm.

Figure 5 illustrates a study of modulated helical electron beam influence distribution for chest wall treatment that was performed using analytical pencil-beam models [64-65]. The helical electron beam energy modulation was generated by superposing several helical electron beams of different energies and weights. The chest wall was treated with at least 90% isodose while the heart received less than 40% isodose. The part of the lung in the region that is closest to the device receives approximately 80% isodose. But the isodose quickly drops to less than 40% as one gets closer to the heart. Thus, usage of this device to treat cancers located on the chest results in lower amount of radiation irradiating the heart and lungs as compared to regular electron beam radiation and compared to other forms of radiation treatment.

Helical electron beams increase the avoidance of radiating healthy tissue during

radiation treatment of cancers. The device helps prevent radiation injury of healthy tissue during radiation therapy of cancers by three different approaches. The first approach is by enhancing the dose of the radiation to a smaller area. The second approach is by lowering the depth and spread of the radiation. And third by allowing one to use smaller areas of treatment than is permitted using standard, known in the art field linear accelerators with a multileaf collimator in the gantry.

All referenced cited herein are incorporated by reference in their entirety.

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While the invention has been described in detail, and with reference to specific embodiments thereof, it will be apparent to one of ordinary skill in the art that various changes and modifications can be made therein without departing from the spirit and scope thereof. The artisan will further acknowledge that the examples recited herein are demonstrative only and are not meant to be limiting.