A linear accelerator for a radiotherapy apparatus

Abstract

A linear accelerator apparatus suitable for radiotherapy applications such as FLASH radiotherapy, comprising an apparatus for accelerating a beam of electrons 412 provided by an injector 410 through t

Description

This disclosure relates to a linear accelerator and in particular but without limitation to a linear accelerator for a radiotherapy apparatus and a method of operating a linear accelerator for a radiotherapy apparatus. Background Radiotherapy can be described as the use of ionising radiation to damage or destroy unhealthy cells in both humans and animals. The ionising radiation may be directed to tumours on the surface of the skin or deep inside the body. Common forms of ionising radiation include X-rays and charged particles. A radiotherapy apparatus or device typically comprises a gantry which supports a beam generation system, or other source of radiation, which is rotatable around a patient. Alternatively, static, fixed beam systems may also be used, sometimes in combination with a rotatable chair or other patient support. The beam generation system is typically based on a particle accelerator such as a linear accelerator or 'linac' which comprises a radiofrequency (RF) power source, a charged particle source, and an RF cavity. Linear accelerators (especially those for medical use) accelerate charged particles such as electrons to relativistic speeds along an acceleration path through an acceleration waveguide. The acceleration waveguide comprises one or more resonant cells that surround the charged particle trajectory. The acceleration waveguide is filled with RF power from an RF power source which forms an oscillating electric field, or an electromagnetic (EM) wave, inside the cavity. Charged particles are injected from a particle source such as an electron gun into the cavity, forming a beam. As the beam traverses the oscillating EM wave, it gains energy and is thus accelerated, often up to relativistic speeds. Accelerated particles with increased effective mass deposit more energy when they collide with other matter, which is usually either biological tissue or a tungsten target, the target being used to generate X-rays. The resultant particle beam, or X-rays, may be used for imaging or treatment, so often the energy of the beam is variable. A very high energy electron (VHEE) beam is a proposed future radiotherapy treatment modality. The high energy of a VHEE beam can enable higher dose delivery depth than conventional electron and x-ray radiotherapy. The very high energy electrons may be used as the treatment beam, rather than converting electrons into x-rays for treatment. The use of very high energies provides the electrons with sufficient penetration depth for the treatment beam to reach deep-seated targets without having been converted to x-rays. VHEE accelerators produce an electron beam with an energy between 120 MeV to 250 MeV that may deliver a dose depth characteristic similar to that of proton therapy machines (which are costly and cannot fit into a standard bunker used for conventional x-ray or electron linear accelerator radiotherapy). Currently, VHEE machines in research facilities are very large and are impractical for medical applications. As such, VHEE is not yet suitable or possible in a commercial or clinical setting. There is therefore a need to develop appropriate VHEE systems for medicine that would be suitable for commercialisation. VHEE technology is also a proposed approach for enabling FLASH radiotherapy (FLASH-RT). FLASH-RT provides an ultra-high dose rate of at least 30 Gy / s at the isocentre. There is a great interest in FLASH-RT research because the potential benefit compared to conventional radiotherapy is reduced toxicity and very short treatment times. Ways to achieve FLASH-RT include high energy heavy particles such as protons (at 250 MeV) and carbon ions, which are highly expensive approaches. X-ray machines using conventional linear accelerator technology have a dose rate that is at least a factor of 10 lower than that required for FLASH-RT, or use electron beams that have insufficient energy to treat deep seated tumours. Compared with an x-ray beam, an electron beam may deliver a dose rate of about 1000 times more than the corresponding dose rate that would be produced by x-rays produced from that same electron beam. An electron beam can thereby provide more efficient energy conversion, making VHEE technology a candidate for FLASH-RT. It is therefore desirable to develop a VHEE system suitable for commercialisation that can enable FLASH-RT. Summary An invention is set out in the independent claims. Optional features are set out in the dependent claims. Examples will now be described, by way of example, with reference to the drawings of which: Fig. 1 shows a radiotherapy device or apparatus; Fig. 2 shows a beam generation system for a radiotherapy device; Fig. 3 shows components of an acceleration waveguide for a linear accelerator; Fig. 4 shows part of a linear accelerator for a radiotherapy apparatus according to the disclosure; Fig. 5 shows a further example of part of a linear accelerator for a radiotherapy apparatus according to the disclosure; Fig. 6 shows a method of operating a linear accelerator for a radiotherapy apparatus according to the disclosure; Fig. 7 shows a block diagram of one implementation of a radiotherapy system; Fig. 8 shows a computer readable medium or, more generally, a computer program product. Overview In overview, a linear accelerator for a radiotherapy apparatus that is capable of producing a VHEE beam is provided. The linear accelerator comprises two or more S-band accelerator sections, each S-band accelerator section coupled to another S-band accelerator section by a corresponding drift tube. A method of operating the linear accelerator for a radiotherapy apparatus is also provided herein, the method comprising providing RF power to the two or more S-band accelerator sections to accelerate a beam of electrons through the two or more S-band accelerator sections. As will be explained in further detail below, such an approach provides a structure with a very high accelerating gradient, producing a VHEE beam in an apparatus that can be sized to fit into conventional linear accelerator radiotherapy apparatus footprints, or can be mounted on a gantry like those of conventional linear accelerator radiotherapy devices. Accordingly, the benefits of VHEE technology can be realised within a linear accelerator sub-system or beam generation system that can, for example, have a length of less than 2.5 m and thus fit on a gantry arm and into a standard radiotherapy bunker space. Such a VHEE beam generation device may be referred to as having a 'space claim' of less than 2.5 m. Furthermore, the apparatus and methods disclosed herein provide a VHEE device that may be less costly and more compact than a proton therapy machine, and may thus enable a new approach to FLASH-RT. Detailed Description Disclosed herein are systems, devices, methods and apparatuses relating to radiotherapy. With linear accelerator-based radiotherapy devices being highly complex and having many inter-related parts, the terms "system", "device", "apparatus", and "machine" may all be applied interchangeably to describe the radiotherapy apparatus as a whole, or collections of components of the radiotherapy apparatus. The term "apparatus" as used herein may refer to either a single apparatus or plural apparatus and should not be understood as being particularly limited to either a single discrete apparatus or a plurality of discrete apparatus unless a particular apparatus is further described as such. Fig. 1 shows an exemplary radiotherapy device 100, or apparatus, suitable for delivering, and configured to deliver, a beam of radiation to a patient during radiotherapy treatment. The radiotherapy system 100 and its constituent components will be described generally for the purpose of providing useful accompanying information for the present disclosure. The radiotherapy system 100 shown in Fig. 1 is suitable for use with the disclosed methods, apparatus, devices, systems, and / or computer readable media. The radiotherapy device 100 is as an image-guided radiotherapy (IGRT) machine. The radiotherapy system 100 comprises a rotatable gantry 102 to which are mounted a treatment apparatus 104 and an imaging apparatus 106. In this example, the treatment apparatus 104 and the imaging apparatus 106 are attached to the gantry, so that they are rotatable with the gantry, i.e. so that they rotate as the gantry rotates. Positioned in a treatment volume 109 of the radiotherapy system 100 is a patient support surface 110 upon which a patient 112 is positioned during radiotherapy treatment. The patient support surface 110 is configured to move between a first position substantially outside the treatment volume 109, and a second position substantially inside the treatment volume 109. In the first position, a patient or subject can mount the patient support surface. The patient support surface 110, and patient, can then be moved inside the bore, to the second position, in order for the patient to be imaged or treated using the radiotherapy system 100. The movement of the patient support surface is effected and controlled by a patient support surface actuator, which may be described as an actuation mechanism. Together, these components may be described as a patient positioning system, which may comprise other components. The patient support surface may also be referred to as a moveable or adjustable couch or table. Treatment apparatus 104 comprises a treatment beam source 114 and a treatment beam target 116. The treatment beam source 114 is configured to emit or direct therapeutic, or treatment, radiation, for example megavolt (MV) energy radiation, towards the treatment volume 109 and thus the patient 112. As the skilled person will appreciate, the treatment beam source 114 may comprise an electron source, a linear accelerator (linac) for accelerating electrons toward a heavy metal, e.g. tungsten, target to produce high energy photons, and a collimator configured to collimate the resulting photons and thus produce a treatment beam. The treatment beam source 114 may accordingly comprise a beam generation system that is arranged to accelerate electrons towards the target to produce a treatment beam. Once the treatment radiation has passed from the treatment beam source 114 and through the patient 112, the treatment radiation continues towards a treatment beam target 116, where it is blocked / absorbed. The treatment beam target 116 may include an imaging panel (not shown). The treatment beam target may therefore form part of an electronic portal imaging device (EPID). EPIDs are generally known to the skilled person and will not be discussed in detail herein. The imaging apparatus 106 comprises an imaging beam source 118 and an imaging panel 120. The imaging beam source 118 is configured to emit or direct imaging radiation, such as X-rays of kV energy, towards the patient 112. As the skilled person will appreciate, the imaging beam source 118 may be an X-ray tube or other suitable source of X-rays. The imaging beam source 119 is configured to produce kV energy radiation. Once the imaging radiation has passed from the imaging beam source 118 and through the patient 112, the imaging radiation continues towards the imaging panel 120. The imaging panel 120 may be described as a radiation detector, or a radiation intensity detector. The imaging panel 120 is configured to produce signals indicative of the intensity of radiation incident on the imaging panel 120. In use, these signals are indicative of the intensity of radiation which has passed through a patient 112. These signals may be processed to form an image of the patient 112. This process may be described as the imaging apparatus 106 and / or the imaging panel 120 capturing an image. By taking images at multiple angles around the patient it is possible to produce a 3D image of the patient, for example using tomographic reconstruction techniques. The imaging beam source 118 may be mounted on an imaging source arm such that the imaging beam source 118 is moveable along a direction parallel to the axis of rotation of the gantry. The imaging source arm is thus configured to deploy the imaging beam source 118 to a position away from the gantry (a deployed position) for use in imaging the patient, and is configured to retract the imaging beam source 118 source to a position near to the gantry (a retracted position) for situations in which imaging is not required. In the illustrated example, the treatment apparatus 104 and the imaging apparatus 106 are mounted on the gantry such that a treatment beam travels in a direction that is generally perpendicular to that of the imaging beam. Because the gantry 102 is rotatable, the treatment beam can be delivered to a patient from a range of angles. Similarly, the patient can be imaged from a range of angles by the imaging apparatus 106. As the skilled person will appreciate, the gantry 102 can be rotated to any of a number of discrete angular positions relative to a patient. The treatment apparatus 104 may direct radiation toward the patient at each or a number of these discrete angular positions, according to a treatment plan. The treatment apparatus 104 may even be used to continuously irradiate a patient at all rotation angles as it is rotated by the gantry 102. The angles from which radiation is applied, and the intensity and shape of the therapeutic beam, may depend on a specific treatment plan pertaining to a given patient. The radiotherapy apparatus 100 additionally comprises a controller (not shown). The controller comprises a computer, processor, and / or other processing device configured to control the radiotherapy apparatus 100. The controller is configured to send control signals to multiple different components of the radiotherapy apparatus 100, for example those described above and elsewhere herein. The controller is also configured to send control signals to the treatment apparatus in order to effect changes in radiotherapy treatment. The controller also collects data indicative of the performance and actions of various components of the radiotherapy apparatus 100. For example, the controller controls rotation of the gantry and records the angle to which the gantry has been rotated. The controller may be formed by several discrete processors; for example, the controller may comprise an imaging apparatus processor, which controls the imaging apparatus 106; an treatment apparatus processor, which controls the operation of the treatment apparatus 104; and a patient support surface processor which controls the operation and actuation of the patient support surface 110. The controller is communicatively coupled to a memory, e.g. a computer readable medium, comprising computer-executable instructions which may be executed by the controller. The computer-executable, or computer-readable, instructions, may cause a processor to perform any one or more of the methods disclosed herein. The radiotherapy device 100 also comprises several other components and systems as will be understood by the skilled person. For example, in order to ensure the linac does not leak radiation, appropriate shielding is also provided. Fig. 2 shows an example beam generation system 200 that will be described generally for the purpose of providing useful accompanying information for the present disclosure. For example, the beam generation system 200 may be used as the beam generation system in the device 100 of Fig. 1. The beam generation system 200 is based on a linear accelerator design. The beam generation system 200 comprises an acceleration waveguide 202 and a source 204 of electrons. The source 204 of electrons may be an electron gun, for example a triode electron gun or diode electron gun. The acceleration waveguide 202 is configured to accelerate particles, in this case electrons, along an acceleration path 206 into a target 208, in order to produce a treatment beam 210 of radiation. The acceleration path 206 is also known as the central beam axis of the acceleration waveguide 202. The acceleration waveguide 202 comprises a series of cells. In this example, each cell has substantially the same shape and dimensions, but in other examples, that may not be so. The cells may be arranged such that each cell is RF-uncoupled / independent, and in that case each cell functions as a separate resonant cavity. In other implementations, such as the example of Fig. 2, the cells may be coupled together and, in that case, the overall coupled structure may be considered to be a single resonant cavity. In such an implementation, although the coupled cells function as a single resonant cavity, individual cells may still be referred to as cavities by those skilled in the art. The acceleration path is coincident with the centre axis of the acceleration waveguide 202 and passes through an aperture at the centre of each cell. The acceleration waveguide 202, the source 204 of electrons, the cells and the target 208 are enclosed within an evacuated and vacuum-sealed casing 212 to ensure that propagation of the electrons is not impeded as they travel toward the target 208. The vacuum-sealed casing 212 is evacuated using a vacuum system to ultra-high vacuum (UHV) conditions. As the electrons are accelerated in the acceleration waveguide 202, in some embodiments the electron beam path may be controlled by a suitable arrangement of steering magnets, or steering coils (not shown), which surround the acceleration waveguide 202. The arrangement of steering magnets may comprise, for example, two sets of quadrupole magnets. A source of RF waves, or RF power source 214, such as a magnetron or a klystron, is configured to produce and / or amplify RF waves. The RF power source 214 is coupled to the acceleration waveguide 202 via an RF transmission apparatus 216, which usually comprises copper waveguide sections that can have a circular or rectangular cross section. A modulator is configured to pulse RF waves through the copper waveguide into the acceleration waveguide 202. Typically, the RF waves are input into a particular cell of the acceleration waveguide 202. The RF transmission apparatus 216 that connects the RF power source 214 to the input cell of the acceleration waveguide 202 may comprise a waveguide network and may contain an RF window which may separate a vacuum envelope from an SF6 envelope. The RF transmission apparatus 216 is perpendicular to the acceleration waveguide 202 central beam axis 206 where it couples the power into the input cell. The RF input connecting pipe ortube is coupled with the acceleration waveguide 202 and joins the acceleration waveguide 202 at a substantially 90° angle. The RF transmission apparatus 216 may include a circulator 218 of any appropriate known type. The beam generation system can operate with either a standing wave or a traveling wave configuration. In a standing wave configuration as shown, the RF power source 214 is configured to pulse RF waves into the acceleration waveguide 202, in order to set up a standing wave of varying electric field that is suitable for accelerating charged particles. Although the RF power source 214 can operate in continuous mode, typically it operates in pulsed mode in view of the RF power levels required. An example RF wave frequency is 3 GHz, with a pulse duration in the range of microseconds and a pulse repetition rate in the range of several hundred pulses per second. The RF power source 214 may be a commercially available magnetron such as an E2V 3.1 MW magnetron, or any standard radiotherapy magnetron, operating at 3 GHz. Typically, the RF power source 214 produces each pulse at a particular phase in order to improve the stability of the standing wave within the acceleration waveguide 202. After it has been pulsed into the acceleration waveguide 202, some of the RF energy dissipates into the walls of the acceleration waveguide 202. In an acceleration waveguide made up of coupled cells, the standing RF wave is established according to the resonant frequency of the coupled structure. An effect of coupling individually resonant cells together to form a single resonant cavity is that, due to dispersion, a band of different frequency oscillation modes comprising higher and lower order modes may be permitted within the acceleration waveguide 202 either side of the resonant frequency of the coupled structure. The frequency of the RF waves provided by the RF power source 214 determines the mode(s) that are excited in the acceleration waveguide 202. There are also multiple modes of operation by which a standing wave at the resonant frequency can accelerate electrons within the acceleration waveguide. Electrons will accelerate or decelerate depending upon the polarity of the electric field they experience. The length of each cell in the cavity is designed such that the beam sees the same phase of the RF in each cell. The beam is synchronised such that on each oscillation the beam interacts with the positive part of the wave and is accelerated further. In one operational mode, known as the zero mode, the electric field of the standing wave has the same polarity and magnitude in all cells at any given time. During the time that an electron takes to traverse a given cell and enter the next cell, the field makes a complete oscillation, for example from positive to negative and back to positive, such that the electron sees the same accelerating field it has just experienced, rather than a decelerating field. Alternatively, a 'n mode' may be used. Rather than the electric field being of the same polarity in each cell at a given time, adjacent cells have opposite polarities at a given time. However, the dimensions of each cell are such that during the time an electron takes to traverse a given cell, the adjacent cell experiences a half oscillation in field polarity such that the electron entering the adjacent cell experiences an accelerating field polarity rather than a decelerating one, and so on. The source 204 of electrons, such as an electron gun, is also coupled to the acceleration waveguide 202 and is configured to inject electrons into the acceleration waveguide 202. The injection of electrons into the acceleration waveguide 202 is synchronised with the pulsing of the radiofrequency waves into the acceleration waveguide 202. In some implementations, an upstream portion of an acceleration waveguide in a linac may be referred to as a buncher section. The buncher section may comprise one or more cells of the acceleration waveguide. Within the buncher section, the phase of the RF wave, whether a standing wave or traveling wave, decelerates some electrons to allow slower electrons to catch up, concentrating the electrons in bunches. The electrons are then free to move together in so called "packets" or "bunches" and the bunches quickly accelerate to relativistic speeds through the subsequent cells of the acceleration waveguide. The acceleration waveguide may be designed with a buncher section that is optimised to produce an electron beam with a particular energy and intensity by bunching electrons into a beam of short pulses. RF waves may be input to the acceleration waveguide at a particular cell, or at more than one cell. In particular, RF waves may be input at a cell that is adjacent to the buncher portion of the acceleration waveguide. In the example of Fig. 2, the first two cells on the left-hand side of the acceleration waveguide 202 are the buncher and the following cells act to accelerate the electrons to relativistic speeds. Alternatively, the RF waves may be input into one or more of the cells belonging to the buncher section of the acceleration waveguide 202. Once the electrons have been accelerated to faster energies, such as 8 MeV or 10 MeV, they may pass into a flight tube. The flight tube is connected to the acceleration waveguide by a connecting tube. The flight tube is also kept under vacuum conditions. This connecting tube or connecting structure is termed a drift tube. The drift tube also forms part of a vacuum tube along with the other components within the vacuum-sealed casing 212. Whilst the electrons travel through the flight tube, an arrangement of focusing magnets act to direct and focus the beam on the target. The electrons travel toward the target 208 which may comprise, for example, tungsten, or another heavy metal. The impact of the electrons on the target 208 produces x-rays which form the treatment beam 210. When the electrons strike the target 208, x-rays are produced in a variety of directions. A primary collimator may block x-rays travelling in certain directions and pass only forward travelling x-rays to produce the treatment beam 210. The x-rays may be filtered and may pass through one or more ion chambers for dose measuring. The beam can be shaped in various ways by beam-shaping apparatus, for example by using a multi-leaf collimator, before it passes into the patient as part of radiotherapy treatment. If a flight tube is used, the target is located inside the flight tube and is located at the end of the flight tube to seal the vacuum system. The flight tube also comprises a target window, which is transparent to x-rays, and which is positioned to allow the x-rays which are produced when the beam generation system is in operation to pass from the evacuated flight tube through the target window and into the treatment head. In some implementations, the electrons are accelerated within an acceleration waveguide by using a travelling wave rather than a standing wave. In this case electrons travel at the phase velocity of the travelling wave, accelerated by the longitudinal electric field component. The acceleration waveguide 202 must be designed such that the phase velocity of the traveling wave does not exceed the speed of light, otherwise no acceleration of electrons will occur. In particular, using a disk-loaded waveguide, rather than a cylindrical waveguide, reduces the phase velocity appropriately such that electrons are accelerated. For an accelerator that uses a traveling wave, in addition to an RF input, the acceleration waveguide will have an RF output configured to transfer RF energy out of the acceleration waveguide and prevent it from reflecting and establishing a standing wave. If a drift tube is used adjacent to the acceleration waveguide, the RF output may be coupled to the drift tube. As with the input transmission apparatus or waveguide, which introduces RF power to the acceleration waveguide, the output waveguide through which RF power exits the waveguide can be connected via an elbow joint or 'T-shaped' joint. RF waves pass out from the evacuated system via an RF output window which seals the vacuum envelope. Fig. 3 shows a cross section view along the longitudinal axis of an acceleration waveguide 300 suitable for use in a particle accelerator, for example for use as the acceleration waveguide 202 of the beam generation system 200 of Fig. 2, and depicts a typical multiple cell cavity. The acceleration waveguide 300 is suitable for use in a linac as shown in Fig. 1 and Fig. 2, but also could be used in other accelerators (e.g. a curved accelerator such as a synchrotron). The below examples and discussion relate to the acceleration of electrons, but the cavity can be used to accelerate any charged particle and therefore in any charged particle accelerator. Two cells 310 of a series of connected cells are shown. The cells may be coupled together. The cells are each connected along a central axis 312 by irises 314, 315. Only two cells are shown in Fig. 3, although a typical acceleration waveguide will have more. The precise number will vary, dependent on the design criteria of the accelerator. Each cell is defined in the form of a recess within a surrounding shell of a conductive material, usually copper. The acceleration waveguide 300 may be described as a disk-loaded waveguide. In the following description, the term "longitudinal cross section" is used to define the cross section in a plane through the centre axis. The "transverse cross section" is used to define the cross section in a plane orthogonal to the centre axis. A longitudinal centre of an object is halfway down the object's longitudinal axis. For example, the longitudinal centre of a cell is the plane half way along the centre axis of that cell. Each cell has an iris 314 connecting to the preceding cell in the sequence, and an iris 315 connecting to the next cell in the sequence. The irises and cells are centred on the centre axis. In use, the centre axis defines the electron acceleration path, the path along which electrons travel when being accelerated though the acceleration waveguide 300. Generally, cells and irises are axisymmetrical around the centre axis, forming a rounded toroid, i.e. the three-dimensional shape created by sweeping a two-dimensional shape around the axis. Although each of the cells shown in Fig. 3 has the same dimensions, with a fixed radius r and fixed length along the acceleration path 312, some waveguide designs may use coupled cells of varying dimensions, in particular cells of varying length. The length of the cell is chosen to provide a suitable length through which each electron experiences acceleration due to the oscillating RF field. In the waveguide shown in Fig. 3, a "nose cone" 316 is formed on each end of the iris, lengthening the iris along the centre axis to protrude into the cell, for purposes of concentrating the longitudinal electric field component at the beam axis. However, some waveguides do not include a nose cone. The cells are manufactured by welding segments of conductive material together at joining portions. The joining portions of the segments are typically in the longitudinal centre of the cell, marked a. Additionally, the plane at the longitudinal centre of each cell is a plane of symmetry. Many alternative methods may be used to make a cavity, such as brazing and diffusion bonding. One alternative to welding individual cavities together is to instead segment a larger cylindrical piece into multiple cavities. Once a waveguide has been produced, the waveguide may be "tuned" to try to bring each cell to the correct resonance. This can be done by taking a measurement of the electric field created in a waveguide upon the application of radiofrequency energy, introducing a perturbation such as a dent into the waveguide, and then taking another measurement of the electric field to determine whether the resonance is correct. This aligns the optimum path the electrons take and minimises their interaction with the higher order modes. Referring to the apparatuses of Fig. 2 and Fig. 3, variations in design and components are possible or desirable depending upon the application requirements. For example, requirements may vary depending upon the desired type and energy of the treatment beam or depending upon the mechanical or structural design of the overall device in which the apparatus is to be used, such as the device 100 of Fig. 1. In some implementations, the RF power source 214 may be a klystron, rather than a magnetron. Similarly, in some implementations, the RF power source 214 may be operated continuously rather than in a pulsed manner. In some implementations, the beam generation system 200 is configured to emit either an x-ray beam or an electron particle beam. Such implementations allow the device to provide electron beam therapy, i.e. a type of external beam therapy where electrons, rather than x-rays, are directed toward the target region as the therapeutic radiation. It is possible to 'swap' between a first mode in which x-rays are emitted and a second mode in which electrons are emitted by adjusting the components of the beam generation system. In essence, it is possible to swap between the first and second mode by swapping between the heavy metal target and a so-called 'electron window'. The electron window may be made of nickel. The electron window is substantially or partially transparent to electrons and allows electrons to exit the beam generation system 200. The beam generation system 200 also comprises several other components and systems as will be understood by the skilled person. For example, in order to ensure the linac or beam generation system does not leak radiation, appropriate shielding is also provided. The whole system is cooled by a water cooling system (not shown in the figures). The water cooling system may be used, in particular, to cool the acceleration waveguide 202, the target 208, and the RF power source 214. Fig. 4 shows part of a linear accelerator 400 for a radiotherapy apparatus. The linear accelerator 400 may be used in the radiotherapy apparatus and systems disclosed herein, such as those of Fig. 1 and Fig. 7. For example, the part of the linear accelerator 400 shown in Fig. 4 maybe used in the treatment beam source 114 of Fig. 1 described above. The part of the linear accelerator 400 shown in Fig. 4 may be used as an acceleration waveguide in a beam generation system. Unlike a conventional radiotherapy acceleration waveguide, which is formed from a continuous series of connected cells or cavities, the linear accelerator 400 has an acceleration waveguide formed of multiple energetically independent, or isolated, or spatially separated, sections. A first section is used as an injector section 410, into which electrons are introduced from a source of electrons, such as those described above, such as an electron gun. Accordingly, an electron beam 412 is formed within the linear accelerator 400. Subsequent sections of the linear accelerator 400 act as consecutive accelerator sections 420a, 420b, 420c to 420n to accelerate the electrons in the electron beam 412 towards the opposite end of the linear accelerator 400. The electron beam 412 may subsequently be output, for example as a beam of therapeutic radiation comprising electrons. The linear accelerator 400 may thus be arranged to produce a beam of therapeutic radiation comprising electrons. Again unlike conventional radiotherapy apparatuses, each section of the linear accelerator 400 acceleration waveguide disclosed herein is isolated from the adjacent section(s) by a drift tube, which provides energetic isolation of sections. For example, the injector section 410 is coupled to a first accelerator section 420a using a drift tube 415, and the further accelerator sections are sequentially coupled using respective drift tubes 425a-c. Although four accelerator sections are shown in Fig. 4, any number n of accelerator sections, including less than four, may be used according to the requirements of the system, so long as each section is isolated from one another and coupled together using a drift tube. In some examples, only two accelerator sections are used. In some implementations, an injector section may be omitted and electrons may be injected into one of the accelerator sections. Each injector and / or accelerator section is made up of one or more RF resonant cavities like those shown in Fig. 3, and those cavities may comprise a side coupling structure. The side coupling structure is arranged to couple a plurality of cavities of the respective section. Each accelerator section 420a-n is formed of one or more cavities and forms an acceleration waveguide structure with a resonant frequency at an S-band RF frequency. Accordingly, the linear accelerator 400 may be arranged to accelerate electrons using an S-band RF frequency. Each accelerator section 420a-n may hence be referred to herein as an S-band accelerator section due to having a resonant frequency in the S-band of radiofrequencies. Each accelerator section 420a-n is arranged to provide the electron beam 412 with relativistic velocity. Each drift tube, on the other hand, is a section which does not accelerate electrons in the electron beam 412, instead allowing the electrons to "drift" without the direct influence of the fundamental accelerating field. The drift tubes 415,425a-c serve to physically connect the sections of the linear accelerator 400 but to isolate them in terms of conduction of RF energy or power. Typically, the drift tubes 415 are arranged to operate at a frequency far below the cut-off frequency associated with the acceleration mode of the adjacent accelerating sections. In particular, each drift tube 415 may be sized with large enough diameter to receive the electron beam 412 but insufficiently large diameter to allow the S-band RF standing wave to pass. In some examples, one or more drift tube sections may have bending magnets provided thereto in order to bend or steer the electron beam 412. Using S-band RF cavities provides a number of advantages. For example, although a small linear accelerator design may conventionally look to use smaller X-band components, the larger cavity size of S-band components allows for a lot of internally stored RF power compared to X-band components. In turn, a high Q factor is provided by S-band components, and in some examples, the Q factor of each accelerator section 420a-n is greater than or equal to 15000. The Q factor is related to the amount of stored power contained within the linac structure, and there is a direct correlation in terms of the on-axis accelerating gradient, measured in MV / m. Therefore, the higher the Q factor, the larger the potential accelerating gradient that can be obtained from a structure. A higher Q factor, as provided by the presently disclosed S-band RF cavities, therefore provides a higher acceleration gradient in terms of MV / m units. Furthermore, in comparison with S-band components, X-band components require a lot more power to be supplied because the X-band component structure does not provide much energy storage, whereas S-band components do not have such high requirements. In some examples, a side-coupled cavity structure, or side coupling structure, is used in one or more accelerator sections 420a-n, further increasing the Q factor of that section and the acceleration gradient provided to the electron beam 412. Accordingly, the linear accelerator 400 of Fig. 4 can achieve acceleration gradients higher than the typical 50-60 MV / m of conventional medical linear accelerators. The linear accelerator 400 thus has a structure that provides a very high accelerating gradient to the electron beam 412 and can be used to obtain a VHEE beam over a relatively short length. In some examples, the linear accelerator 400 need only be provided with 19 MW of peak RF power to the S-band accelerator sections in order to accelerate the beam of electrons to at least 120 MeV. Advantageously, this peak RF power required can be 2 to 3 times lower than required in, for example, X-band systems, which may also need to use pulse compression techniques to handle their associated higher energies. Further advantageously, a commercially available S-band klystron may be used to provide 19 MW of peak RF power to the linear accelerator 400. In some examples, the linear accelerator 400 may be mounted to a rotatable gantry and is thus arranged to be mounted to or on a rotatable gantry, or is mountable to or on a rotatable gantry. The rotatable gantry may be a conventional radiotherapy treatment gantry such as that of Fig. 1. Charged particles injected into an acceleration waveguide can form a beam with a particular current and can produce effects due to "beamloading" within the acceleration waveguide based on the amount of that current. Generally, there are three "beamloading" states or operating conditions in which a linear accelerator could be operated: under-beamloaded, beamloaded, and over-beamloaded. Each beamloading operating condition represents a different amount and / or level of beamloading of the acceleration waveguide. Each of the three different beamloading operating conditions can produce distinctly different characteristics in terms of the dose rate, effect on the RF power source, and the RF tune of the structure. Conventional medical linear accelerator beam generation systems are primarily designed to be operated in a "beamloaded" operating condition to obtain a maximum dose rate, and may rely on the effects of a sufficient amount and / or level of beamloading to enable a desired dose to be generated. In general, beamloading occurs when the particle beam (which may be any charged particle such as electrons, protons or carbon ions) absorbs energy from the RF power source to be accelerated to a higher energy. This lowers the amount of stored energy within the linac and therefore lowers the acceleration gradient. Conventional medical linacs are designed and operated in a "beamloaded" operating condition, lying between being under-beamloaded and over-beamloaded, that is referred to as being "peaked", which is a condition in which a maximum dose rate condition is obtained, controlled by an automatic frequency control system. In this state, the RF power source is directly affected by the beam, as is the RF tuning of the structure, which can be detuned as a result, and the beam generation system as a whole exhibits emergent system behaviour as a result of this interaction. The result is a highly non-linear relationship between the electromagnetic fields within the structure and the dose rate produced. Any further additional beam current introduced into the linac past this "peaked" point will have a detrimental effect on the system as the linac has a defined fill time (a period in which the RF power must be supplied to produce electromagnetic accelerating field). Once the particle beam absorbs more energy than can be supplied to the linac, the dose rate will begin to decrease exponentially as a result of both the linac being detuned and the lack of acceleration field from the RF power source. As a result, the system enters the "over-beamloaded" regime, which is conventionally not desirable and is not used as a mode of operation in medical linacs as a result. An "under-beamloaded" operating condition is a state in which the beam current is low enough that the RF power source is unaffected by the particle beam being accelerated. This may mean that the electromagnetic fields within acceleration waveguide behave in a linear or relatively linear manner and the particle beam does not cause detuning of the RF power source and / or acceleration waveguide. In a conventional medical linac, the dose rate of an "under-beamloaded" operating condition is considered suboptimal and so conventional medical linacs do not operate in an under-beamloaded condition for delivering treatment. The present inventors have identified that the under-beamloaded operating condition of linear accelerators such as those of Fig. 4 can provide several benefits, such as offering a very linear regime to operate within and / or enabling the energy gradient per unit length of the acceleration waveguide to be increased by correspondingly accepting or providing a decreased beam current. In a conventional photon or x-ray-based linear accelerator, such a decrease in beam current would typically render such a machine unusable, but in the present disclosure, may be used to facilitate VHEE beam production. The present inventors have identified that a VHEE beam may be produced by a level of beamloading that is different to that used in conventional x-ray-based linear accelerators and is also different to that used in conventional electron therapy linear accelerators. In particular, the level of beamloading used to produce a VHEE beam may fall between the level of a conventional electron therapy linear accelerator and the level of a conventional x-ray-based linear accelerator. As will be understood by those skilled in the art, an "under-beamloaded", "beamloaded", or "over-beamloaded" condition of a particular system will be defined by the specific design of the system, and the level of beamloading may generally be represented by the beam current obtained at the target for a given input RF energy and input particle current. In general, a beam may be operated between open circuit and fully loaded, depending on the machine. It will be appreciated that various approaches, as will be known to the skilled person, may be used to determine an amount or level of beamloading for a specific system, such as the approaches described in Beam Loading In Linear Accelerators by J. E. Leiss, National Bureau of Standards, IEEE Transactions on Nuclear Science, 1965. A non-limiting illustrative example will now be described in order to demonstrate that the linear accelerator 400 of Fig. 4 may enable a VHEE beam to be produced over a distance small enough that the apparatus may fit into a standard linac radiotherapy space, and / or the linear accelerator may be mountable on a rotatable gantry. It will be understood that the exact dimensions and parameter values described are non-limiting and that alternative dimensions and values of operating parameters may be used for different applications and requirements. The present inventors have identified that, in one example, a reduction in beamloading of between 80% to 90% for a high Q (15000 or more) structure can result in a beneficial increase in energy gradient within a 2.5 m length. Accordingly, an energy range between 120MeV to 180MeV with a dose rate at an electron window of at least 26kGy / s can be achieved. In particular, this may be achieved using a design in which an injector section is used with four S-band accelerator sections, with each accelerator section designed to take 3 MW of RF power. Such a structure has a length of around 2.5 m, suitable for mounting on a gantry and / or placing in a conventional linac radiotherapy bunker. In a typical linac radiotherapy device operated with intentional beamloading, an electron beam corresponding to 500 mA is input into the accelerating waveguide, with around 1 MW of RF power taken into the electron beam for acceleration of the electrons. However, for VHEE acceleration in the present standing wave linear accelerator 400 structure, such high beam current is not required because the electrons themselves are used as a beam of therapeutic radiation and therefore the electrons do not need to hit a target in sufficient quantity to provide a suitable X-ray dose as in conventional linac radiotherapy. As a consequence, high electron energies can be achieved by providing the linear accelerator 400 with a 50 mA electron beam, thereby operating in an under-beamloaded condition that enables those electrons to be accelerated at high gradient with relatively low power consumption. Accordingly, applying 21 MW of input power to the linear accelerator 400, along with 50 mA - 100 mA of input electron beam current, can provide a 120 MeV - 180 MeV energy electron beam over 2.5 m of accelerator, because of the approximately 90 MV / m acceleration gradient provided by the S-band isolated section structure described herein. Alternatively, if operated in fully beamloaded condition with 500mA of input from the source of electrons or electron gun, the then output energy of the structure will be within a conventional treatment range of 6 MeV -25 MeV. In general, an under-beamloaded condition for the linear accelerators disclosed herein may be taken to mean that less than 500 mA of electron input is provided for acceleration by a source of electrons such as an electron gun. In some examples, the under-beamloaded condition may be taken to mean that 50 mA -100 mA of electron input is provided for acceleration by a source of electrons such as an electron gun. As will be understood by the skilled person, in some examples, appropriate bending and / or scanning magnets may be used with the VHEE linear accelerators disclosed herein, such as the linear accelerator 400 of Fig. 4. For example, a single pole alpha magnet may be used as a bending magnet, since the VHEE beam is a monochromatic / monoenergetic beam. A scanning magnet may, for example, enable a preferential dose distribution. Furthermore, appropriate use of bending and / or scanning magnets may advantageously enable the overall size of the device to be reduced. Further illustrative examples of VHEE capability that may be provided by a linear accelerator in accordance with that of Fig. 4 are provided in Table 1. It will again be understood that the exact dimensions and parameter values described are non-limiting and that alternative dimensions and values of operating parameters may be used for different applications and requirements. For example, in some examples, a VHEE beam with energy of 200 MeV, or in the range of 180 MeV to 250 MeV, may be produced. Example number Approximate device length Electron beam energy Electron input (gun) current RF power source Dose rate at electron window 1 0.4 m 6 MeV to 7 MeV 500 mA 3 MW at 275 pulse repetition frequency (PRF) and 5 microsecond pulse 130 kGy / s 2 2.5 m 120 MeV to 180 MeV 50 mA to 100 mA 19 MW at 275 PRF and 5 microsecond pulse 26 kGy / s 3 5.4 m 250 MeV to 300 MeV 5 mA to 10 mA 38 MW at 275 PRF and 5 microsecond pulse 2.6 kGy / s Table 1 A particular advantage in having the accelerator sections separated by drift tube sections is that the 5 drift tubes provide "isolation" of each accelerator section in terms of the response of each accelerator section to being supplied with RF power. The accelerator sections can therefore be powered in different ways to provide different energy and dose modulation which would not be available if using a conventional "monoblock" series of non-isolated cavities. Accordingly, each of the S-band accelerator sections is arranged to be independently powered with a respective amount of RF power. 10 In some examples, each accelerator section is isolated from the next by approximately 5 cm of distance, and the first accelerator section is isolated from the injector section with approximately 5 cm of distance. In one example, the linear accelerator 400 is powered by a klystron (in one example, a single pulsed 15 klystron is used to provide 21 MW of pulse power), and very little is required in terms of a complex control system as all the accelerating sections 420a-n are of the same frequency. Accordingly, the klystron is arranged to provide RF power to the S-band accelerator sections 420a-n. In another example, a plurality of pulsed magnetrons, which may be in the commercially available 3 MW to 8 MW range, are used to either power the individual sections or, by using a series of 3 dBm hybrid combiners, to power cojoined sections in order to provide the accelerating gradient. The use of multiple magnetrons requires a more complex phase control (locking RF system) and the greater the number of magnetrons the greater the complexity of this system becomes. However, the multiple magnetron power example lends itself to pulse interleaving whereby it would be possible to generate different energies pulse to pulse by varying the RF power in the isolated accelerator sections 420a-n. Pulse interleaving also allows for dose variation if desired. A similar effect could also be achieved with a single source klystron and 3 dBm combiner connected to each section to which the input power to each section would be controlled by electromechanical means into a load. Energy and dose variation may also be achieved, as in a conventional radiotherapy linac, by varying the electron gun current and from RF power source variation. The linear accelerator 400 is thus arranged to accelerate the beam of electrons by beamloading each S-band accelerator section equally and modulating the source of electrons. In one example, each accelerator section is beamloaded equally and an electron gun input is used to modulate the energy. In another example, the on-axis electrical field of each accelerator section may be varied using an approach based on cell shorting with side-coupled energy selection, such as the approach described in Elekta Limited UK patent application 2016205.3. In order to deliver the VHEE beam produced by the linear accelerator 400 to a patient in the form of a beam of therapeutic radiation, a thin nickel window may be provided at the end of linear accelerator 400 to function as an electron window. An electron beam of, for example, 50mA to 100mA can then be steered to the patient using a bending system to a suitable delivery system. The delivery system may be similar to that used conventionally, such as a rastering pencil beam steering system, a multileaf collimator designed for electrons, or a scattering applicator. In one example, the linear accelerator 400 may comprise a photon target at the end, and may have a photon window. Such an example may enable to production of X-rays for treatment. In one example, a switchable electron / photon window component may be used, such that the linear accelerator 400 may alternate between electron and photon treatment modalities. The linear accelerator 400 is particularly suited to such an application, because an under-beamloading operating condition may be used to produce a VHEE beam, whereas a beamloaded operating condition may be used to produce electrons suitable for hitting a target and producing X-rays, as described above. The sections of linear accelerator 400 may each be joined together using a respective suitable joint, such as an RF component joint suitable for ultra-high vacuum, such as that described in Elekta UK patent application number 2007977.8. In some examples, the joint(s) may comprise a flange, such as a CF (ConFlat) flange with a gasket placed between two knife-edges, or a CF choke flange with a choke groove. In other examples, an alternative conventional "knife edge" connection may be used. Each connecting interface with each drift tube may be matched in phase length and energy, such that the longitudinal and transverse emittance may be matched between sections. Fig. 5 shows a linear accelerator 500 which comprises the linear accelerator 400 of Fig. 4 and further includes one or more quadrupole magnets 512, 514a-c. Each quadrupole magnet 512, 514a-c is arranged to act upon a corresponding drift tube 415, 425a-c, and may act to focus the electron beam within the drift tube. It will be understood that, in some examples, not every drift tube may have a corresponding quadrupole magnet, and indeed, in some examples, only one quadrupole magnet unit will be present in the linear accelerator, which may be arranged to act upon any one of the drift tubes. Although quadrupole magnets 512, 514a-c are preferably used, in some examples, dipole magnets may be used in place of any of the instances of quadrupole magnets 512, 514a-c disclosed herein. Each quadrupole magnet 512, 514a-c provides a further modulation possibility to the linear accelerator 500. For example, in addition to or instead of the energy modulation approaches set out above, the energy of the electron beam 412 may be modulated between respective accelerator sections using the corresponding quadrupole magnet, which may give dose depth distribution of the resultant treatment beam. The linear accelerator 500 is therefore arranged to accelerate the beam of electrons by adjusting the energy of the beam of electrons using the one or more quadrupole magnets 512, 514a-c. Additionally or alternatively, one or more of the quadrupole magnets 512, 514a-c may be arranged to smear the electron beam 412 by smearing bunches of electrons within one or more of the drift tubes of the linear accelerator. Smearing the electron beam 412 may provide an elliptical shaped beam for treatment purposes. Smearing, as used herein, means to shape the electron beam cross-section to deviate from an idealised circular cross-section geometry to, for example, an elliptical cross-section. Further additionally or alternatively, one or more of the quadrupole magnets may be used to raster the electron beam 412 and tailor its shape for treatment purposes. For example, a microbeam effective treatment may be implemented using smearing. Fig. 6 shows a method 600 of operating the VHEE linear accelerators 400, 500 disclosed herein. At a first step 610, the method comprises providing RF power to the two or more S-band accelerator sections to accelerate a beam of electrons through the two or more S-band accelerator sections. The RF power may be provided to each accelerator section. The RF power may be provided by an RF source such as the klystrons and magnetrons disclosed herein. At a second step 620, the method 600 optionally comprises providing a respective amount of RF power to each of the two or more S-band accelerator sections independently. Such RF power may be modulated or supplied independently according to any one or more of the multiple approaches set out above in relation to Fig. 4, such as by using multiple respective power sources. The method 600 may further comprise the steps already set out in relation to the operation of the linear accelerators 400, 500 of Figs. 4 and 5. For example, the method 600 may further comprise producing a beam of therapeutic radiation comprising electrons. In some examples, the method 600 further comprises providing the beam of electrons such that the linear accelerator operates in an under-beamloaded condition. In one example, an under-beamloaded condition corresponds to a beam of electrons of 50 to 100 mA. In another example, an under-beamloaded condition corresponds to a beam of electrons of less than 500 mA. Likewise, the method 600 may further comprise accelerating the beam of electrons to an energy between 120 MeV and 180 MeV using an under-beamloaded operating condition. The method 600 may further comprise accelerating the beam of electrons by beamloading each S-band accelerator section equally and modulating the source of electrons. If the method 600 is used to operate the linear accelerator 500 of Fig. 5, the linear accelerator will further comprise one or more quadrupole magnets, each quadrupole magnet corresponding to a respective drift tube, and the method 600 may further comprise accelerating the beam of electrons by adjusting the energy of the beam of electrons using the one or more quadrupole magnets. Furthermore, the method 600 may further comprise smearing bunches of electrons within one or more of the drift tubes using the one or more quadrupole magnets. The apparatuses and methods disclosed herein, such as those of Figs. 4 to 6, may provide a VHEE system that in turn can enable flash radiotherapy (FLASH-RT). Accordingly, disclosed herein is a flash radiotherapy apparatus comprising the linear accelerator of Fig. 4 and / or 5. VHEE may provide advantages over x-ray-based approaches to FLASH-RT, because energy is lost in converting electrons to x-rays such that x-ray-based approaches require very high power acceleration of electrons to achieve high dose rates, whereas a VHEE based approach does not lose energy in that conversion step. In order to deliver FLASH-RT, a suitable treatment beam delivery system is required. However, for FLASH-RT to be effective in a single fraction, doses of around 30 Gy / s are required, which is difficult to support with a rotating gantry. Accordingly, conventional rotating gantry configurations such as the one shown in Fig. 1 may be modified to meet this new requirement. For example, conventional systems may be modified to allow for a separate FLASH-RT treatment modality to operate alongside a VHEE treatment modality. In one example, a magnetic steering system may be used to steer the VHEE electron beam to a separate target or targets for FLASH-RT. In such examples, a beam steering system may be implemented around the gantry to direct the electron beam to a series of electron targets, and / or a continuous electron window that runs along the inside of the gantry. The rotation of the gantry is inhibited during this treatment modality and the treatment beam is steered to the target / targets to allow for a single fraction treatment to occur. In another example, the gantry may use a very fast drive system, capable of at least 200 RPM, with provision of tomotherapy-style treatment. In such examples, the mechanism for rotating the gantry to allow for the standard VHEE treatment may be temporarily disabled and switched to a more powerful drive system (for example, a motor, pneumatics, or a high pressure driven system). A pencil beam delivery system using a VHEE beam in a similar manner to an x-ray tomotherapy style system may be employed with the gantry rotating at at least 200 RPM to deliver high dose rates. In such approaches, real-time CT imaging may be used to provide rapid imaging, providing very high precision. In another example, varying the treatment energy (via beam loading effects and energy variation mechanisms disclosed herein) along with a pencil beam or microbeam arrangement may produce a combination of electron bunches to allow for variation of treatment depth and provide a three-dimensional dose profile. Using a microbeam generated from smearing of the electron beam (as previously outlined herein) allows for the treatment beam to be shaped to match the cross-section of the tumour at a given depth, while the effective treatment depth may be controlled via energy modulation of the VHEE beam, producing an equivalent to a 3D treatment. This example may either be implemented with the gantry in a fixed position or with a very fast gantry rotation of at least 200 RPM. A computer-based system may be used for controlling or operating various parts of the systems, devices, methods and apparatuses disclosed herein. The computer-based system can be implemented in software, firmware and / or hardware and may comprise a computer-readable medium containing instructions that, when executed by a processor, cause the system to perform any of the methods described herein. Figure 7 illustrates a block diagram of one implementation of a radiotherapy system 700. The radiotherapy system 700 comprises a computing system 710 within which a set of instructions, for causing the computing system 710 to perform any one or more of the methods discussed herein, may be executed. The computing system 710 shall be taken to include any number or collection of machines, e.g. computing device(s), that individually or jointly execute a set (or multiple sets) of instructions to perform any one or more of the methods discussed herein. That is, hardware and / or software may be provided in a single computing device, or distributed across a plurality of computing devices in the computing system. In some implementations, one or more elements of the computing system may be connected (e.g., networked) to other machines, for example in a Local Area Network (LAN), an intranet, an extranet, or the Internet. One or more elements of the computing system may operate in the capacity of a server or a client machine in a client-server network environment, or as a peer machine in a peer-to-peer (or distributed) network environment. One or more elements of the computing system may be a personal computer (PC), a tablet computer, a set-top box (STB), a Personal Digital Assistant (PDA), a cellular telephone, a web appliance, a server, a network router, switch or bridge, or any machine capable of executing a set of instructions (sequential or otherwise) that specify actions to be taken by that machine. The computing system 710 includes controller circuitry 711 and a memory 713 (e.g., read-only memory (ROM), flash memory, dynamic random access memory (DRAM) such as synchronous DRAM (SDRAM) or Rambus DRAM (RDRAM), etc.). The memory 713 may comprise a static memory (e.g., flash memory, static random access memory (SRAM), etc.), and / or a secondary memory (e.g., a data storage device), which communicate with each other via a bus (not shown). Controller circuitry 711 represents one or more general-purpose processors such as a microprocessor, central processing unit, accelerated processing units, or the like. More particularly, the controller circuitry 711 may comprise a complex instruction set computing (CISC) microprocessor, reduced instruction set computing (RISC) microprocessor, very long instruction word (VLIW) microprocessor, processor implementing other instruction sets, or processors implementing a combination of instruction sets. Controller circuitry 711 may also include one or more special-purpose processing devices such as an application specific integrated circuit (ASIC), a field programmable gate array (FPGA), a digital signal processor (DSP), network processor, or the like. One or more processors of the controller circuitry may have a multicore design. Controller circuitry 711 is configured to execute the processing logic for performing the operations and steps discussed herein. The computing system 710 may further include a network interface circuitry 718. The computing system 710 may be communicatively coupled to an input device 720 and / or an output device 730, via input / output circuitry 717. In some implementations, the input device 720 and / or the output device 730 may be elements of the computing system 710. The input device 720 may include an alphanumeric input device (e.g., a keyboard or touchscreen), a cursor control device (e.g., a mouse or touchscreen), an audio device such as a microphone, and / or a haptic input device. The output device 730 may include an audio device such as a speaker, a video display unit (e.g., a liquid crystal display (LCD) or a cathode ray tube (CRT)), and / or a haptic output device. In some implementations, the input device 720 and the output device 730 may be provided as a single device, or as separate devices. In some implementations, the computing system 710 may comprise image processing circuitry 719. Image processing circuitry 719 may be configured to process image data 780 (e.g. images, or imaging data), such as medical images obtained from one or more imaging data sources, a treatment device 750 and / or an image acquisition device 740. Image processing circuitry 719 may be configured to process, or pre-process, image data. For example, image processing circuitry 719 may convert received image data into a particular format, size, resolution or the like. In some implementations, image processing circuitry 719 may be combined with controller circuitry 711. In some implementations, the radiotherapy system 700 may further comprise an image acquisition device 740 and / or a treatment device 750, such as those disclosed herein in the example of Fig. 1. The image acquisition device 740 and the treatment device 750 may be provided as a single device. In some implementations, treatment device 750 is configured to perform imaging, for example in addition to providing treatment and / or during treatment. The treatment device 750 comprises the main radiation delivery components of the radiotherapy system, such as the beam generation systems and linear accelerator components disclosed herein. Image acquisition device 740 may be configured to perform positron emission tomography (PET), computed tomography (CT), and magnetic resonance imaging (MRI). Image acquisition device 740 may be configured to output image data 780, which may be accessed by computing system 710. Treatment device 750 may be configured to output treatment data 760, which may be accessed by computing system 710. Computing system 710 may be configured to access or obtain treatment data 760, planning data 770 and / or image data 780. Treatment data 760 may be obtained from an internal data source (e.g. from memory 713) or from an external data source, such as treatment device 750 or an external database. Planning data 770 may be obtained from memory 713 and / or from an external source, such as a planning database. Planning data 770 may comprise information obtained from one or more of the image acquisition device 740 and the treatment device 750. The various methods described above may be implemented by a computer program. The computer program may include computer code (e.g. instructions) 810 arranged to instruct a computer to perform the functions of one or more of the various methods described above. The steps of the methods described above may be performed in any suitable order. For example, step 620 of method 600 may be performed before, after, simultaneously or substantially simultaneously with step 610. The computer program and / or the code 810 for performing such methods may be provided to an apparatus, such as a computer, on one or more computer readable media or, more generally, a computer program product 800)), depicted in Figure 8. The computer readable media may be transitory or non-transitory. The one or more computer readable media 800 could be, for example, an electronic, magnetic, optical, electromagnetic, infrared, or semiconductor system, or a propagation medium for data transmission, for example for downloading the code over the Internet. Alternatively, the one or more computer readable media could take the form of one or more physical computer readable media such as semiconductor or solid state memory, magnetic tape, a removable computer diskette, a random access memory (RAM), a read-only memory (ROM), a rigid magnetic disc, and an optical disk, such as a CD-ROM, CD-R / W or DVD. The instructions 810 may also reside, completely or at least partially, within the memory 713 and / or within the controller circuitry 711 during execution thereof by the computing system 710, the memory 713 and the controller circuitry 711 also constituting computer-readable storage media. In an implementation, the modules, components and other features described herein can be implemented as discrete components or integrated in the functionality of hardware components such as ASICS, FPGAs, DSPs or similar devices. A "hardware component" is a tangible (e.g., non-transitory) physical component (e.g., a set of one or more processors) capable of performing certain operations and may be configured or arranged in a certain physical manner. A hardware component may include dedicated circuitry or logic that is permanently configured to perform certain operations. A hardware component may comprise a special-purpose processor, such as an FPGA or an ASIC. A hardware component may also include programmable logic or circuitry that is temporarily configured by software to perform certain operations. In addition, the modules and components can be implemented as firmware or functional circuitry within hardware devices. Further, the modules and components can be implemented in any combination of hardware devices and software components, or only in software (e.g., code stored or otherwise embodied in a machine-readable medium or in a transmission medium). It is to be understood that the above description is intended to be illustrative, and not restrictive. Many other implementations will be apparent to those of skill in the art upon reading and understanding the above description. Although the present disclosure has been described with reference to specific example implementations, it will be recognized that the disclosure is not limited to the implementations described, but can be practiced with modification and alteration within the spirit and scope of the appended claims. Accordingly, the specification and drawings are to be regarded in an illustrative sense rather than a restrictive sense. The scope of the disclosure should, therefore, be determined with reference to the appended claims, along with the full scope of equivalents to which such claims are entitled.

Claims

19 03 251. A linear accelerator for a radiotherapy apparatus, the linear accelerator comprising:two or more S-band accelerator sections, each S-band accelerator section coupled to another S-band accelerator section by a corresponding drift tube,wherein the linear accelerator is arranged to accelerate a beam of electrons through the two or more S-band accelerator sections, wherein the linear accelerator further comprises a source of electrons arranged to produce the beam of electrons, and the linear accelerator is arranged to accelerate the beam of electrons by:beamloading each S-band accelerator section substantially equally; andmodulating the source of electrons..

2. The linear accelerator of claim 1, wherein the linear accelerator is arranged to produce a beam of therapeutic radiation comprising electrons.

3. The linear accelerator of claim 1 or claim 2, wherein each S-band accelerator section includes a side coupling structure arranged to couple a plurality of cavities of the respective S-band accelerator section.

4. The linear accelerator of any preceding claim, further comprising an injector section that is coupled to a first of the two or more S-band accelerator sections by a drift tube.

5. The linear accelerator of any preceding claim, wherein the linear accelerator is arranged to be mounted on a rotatable gantry.

6. The linear accelerator of any preceding claim, further comprising a radio frequency, RF, power source arranged to provide at least 19 MW of pulsed RF power to the S-band accelerator sections in order to accelerate the beam of electrons to at least 120 MeV.19 03 257. The linear accelerator of any preceding claim, wherein each S-band accelerator section has a Q factor of at least 15000.

8. The linear accelerator of any preceding claim, wherein each of the S-band accelerator sections is arranged to be independently powered with a respective amount of RF power.

9. The linear accelerator of any preceding claim, further comprising a klystron arranged to provide RF power to the two or more S-band accelerator sections.

10. The linear accelerator of any preceding claim, wherein the linear accelerator is arranged to accelerate the beam of electrons to an energy of between 120 MeV and 180 MeV using an under-beamloaded operating condition.

11. The linear accelerator of any preceding claim, wherein each S-band accelerator section is arranged to provide an accelerating gradient of at least 90 MV / m to the beam of electrons.

12. The linear accelerator of any preceding claim, further comprising one or more quadrupole magnets, each quadrupole magnet corresponding to a respective drift tube, optionally wherein the linear accelerator is arranged to accelerate the beam of electrons by adjusting the energy of the beam of electrons using the one or more quadrupole magnets.

13. The linear accelerator of claim 12, wherein the linear accelerator is arranged to smear bunches of electrons within a drift tube using the one or more quadrupole magnets.

14. A method of operating the linear accelerator of any of claims 1 to 13, the method comprising providing RF power to the two or more S-band accelerator sections to accelerate a beam of electrons through the two or more S-band accelerator sections.

15. The method of claim 14, the method further comprising producing a beam of therapeutic radiation comprising electrons.19 03 2516. The method of claim 14 or claim 15, wherein providing RF power to the two or more S-band accelerator sections comprises providing a respective amount of RF power to each of the two or more S-band accelerator sections independently.

17. The method of any of claims 14 to 16, the method further comprising providing the beam of electrons such that the linear accelerator operates in an under-beamloaded condition.

18. The method of any of claims 14 to 17, the method further comprising accelerating the beam of electrons to an energy of between 120 MeV and 180 MeV using an under-beamloaded operating condition.

19. The method of any of claims 14 to 18, the method further comprising accelerating the beam of electrons by adjusting the energy of the beam of electrons using one or more quadrupole magnets.

20. The method of any of claims 14 to 19, the method further comprising smearing bunches of electrons within a drift tube using a quadrupole magnet.

21. A computer-readable medium containing instructions that, when executed by a processor, cause the performance of the method of any of claims 14 to 20.

22. A radiotherapy apparatus comprising the linear accelerator of any of claims 1 to 13.

23. A flash radiotherapy apparatus comprising the linear accelerator of any of claims 1 to 13.

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