GB2424281A - Radiotherapeutic Apparatus with MRI - Google Patents

Abstract

A radiotherapeutic apparatus comprises a linear accelerator that is rotateable around an isocentre, and a magnetic resonance imaging apparatus disposed around the isocentre, in which the rf coil windings of the magnetic resonance imaging system are of a conductive material shaped as a thin strip of the same order as the skin effect of radio-frequency signal conduction. A significantly better path for transmission of radiation to the patient is thus obtained, with less inhomogeneity and scattering. The rf coil windings preferably have a thickness less than about 100 microns, about 15 times the theoretical skin depth. Further, a radiotherapeutic apparatus comprises a linear accelerator that is rotateable around an isocentre within a beam transmission plane, and a magnetic resonance imaging apparatus disposed around the isocentre, the magnetic resonance imaging apparatus including coils for generating a magnetic field, disposed in an annular ring located in a plane displaced from the beam transmission plane. Placing the coils of the main magnet and/or the gradient coils on either side of the beam portal provides what are, in effect, annular shielding elements either side of the beam. This means that radiation scattered from parts of the MRI system can be absorbed by the coils in their capacity as shields, and (therefore) that the amount of further shielding can be reduced or even eliminated.

Description

Radiotherapeutic apparatus

FIELD OF THE INVENTION

The present invention relates to a radiotherapeutic apparatus. It seeks to include the functions of magnetic resonance imaging (MRI) in a radiotherapeutic apparatus and to overcome the difficulties presented in doing so.

SUMMARY OF THE INVENTION

In its first aspect, the present invention therefore provides a radiotherapeutic apparatus comprising a linear accelerator that is rotateable around an isocentre, and a magnetic resonance imaging apparatus disposed around the isocentre, in which the rf coil windings of the magnetic resonance imaging system are of a conductive material shaped as a strip of thickness not greater than dmax where: d11=15 / 2 V /i0c7ü) where Po is the magnetic permeability of vacuum, a is the electrical conductivity and w is the frequency times 2it.

Given that magnetic permeability of vacuum is 4ir107 in Hm', the electrical conductivity is S/m (or A2s3/kgm3 and the frequency is Hz, this will yield a value for the skin depth in meters.

We have found that coil windings of this type offer a significantly better path for transmission of radiation to the patient, with less inhomogeneity and scattering. This is however achieved with little impact on the properties of the rf coil, since electrical signals of this frequency are typically carried in only the surface layers of the conductor - the so-called "skin effect". Since the above limit will, in extremis, allow the conductor still to be about 15 times the depth of the skin, the effect on the electrical properties of the coil is likely to be minimal.

The rf coil windings preferably have a thickness less than about 100 microns, which in the case of copper is about 15 times the theoretical skin depth given by the above equation. They preferably have a thickness less than about microns, however, since further reduction should assist their transmission properties still further. Indeed, it is generally preferred that the thickness of the rf coil windings is less than: d=5/ 2

I

which in practice will be satisfied by a thickness of about 30 microns.

It is preferred that the rf coil windings are mounted on a non-metallic carrier, such as PerspexTM.

In a second aspect, the present invention also relates to a radiotherapeutic apparatus comprising a linear accelerator that is rotateable around an isocentre within a beam transmission plane, and a magnetic resonance imaging apparatus disposed around the isocentre, the magnetic resonance imaging apparatus including coils for generating a magnetic field, disposed in an annular ring located in a plane displaced from the beam transmission plane.

Generally, it is preferable for the coil system to be symmetrical, and thus the coils of the magnetic resonance imaging apparatus can be placed in a first annular ring displaced from the beam transmission plane in a first direction and also in a second annular ring displaced from the beam transmission plane in a second direction opposite to the first direction.

Placing the coils of the main magnet and/or the gradient coils on either side of the beam portal provides what are, in effect, annular shielding elements either side of the beam. This means that radiation scattered from parts of the MRI system can be absorbed by the coils in their capacity as shields, and (therefore) that the amount of further shielding can be reduced or even eliminated.

BRIEF DESCRIPTION OF THE DRAWINGS

An embodiment of the present invention will now be described by way of example, with reference to the accompanying figures in which; Figure 1 shows a view of the device according to the invention.

Figures 2a and 2b show iso-gradient profiles of the magnetic field with a cm gap in the gradient coils. Figure 2a shows the transversal mid-plane in which the gradient quality is similar to a conventional (closed) gradient coil.

Figure 2b shows the sagital mid-plane, in this direction the field of view is limited to 25 cm. The distances on the axes are in meters.

Figures 3a and 3b show a schematic view of the layered structures within a MRI system and the scatter induction as function of the field size. Figure 3b shows the schematic of the layered structures within a MRI system, and Figure 3b shows scatter induction as function of the field size, measured for a slab of 10 cm aluminium and an internal wedge and calculated for the layered configuration as shown in figure 3(a). The scatter induction is determined at 10 cm from the

field edge.

Figure 4 shows a schematic drawing of the arrangement used for the MonteCarlo simulations of figures 5a and 5b.

Figures 5a and 5b show the results of Monte Carlo simulations showing the cross-section for a 6 MV pencil beam with the presence and absence of a 1. 5 T magnetic field at 1 cm depth. The magnetic field is oriented from bottom to top in the plane of paper, and the plot is normalised on 10% of the maximum dose in order to show the induced asymmetry more clearly. Figure 5a shows the result without the applied magnetic field, and figure 5b shows the result with the

1.5T magnetic field.

Figures 6a and 6b show calculated central dose profiles in the X direction at 5cm depth and the central axis depth-dose curves for a lxi and a 5x5 cm2 field (solid lines and dashed lines, respectively), both with and without a 1.5T magnetic field (thick and thin lines, respectively). Figure 6a shows the lateral does profiles at 5cm depth, and figure 6b shows the depth-dose curves.

Figure 7 shows a schematic illustration of an air slit geometry.

Figures 8a and 8b show the dose profile on a central axis in comparison to the situation without a magnetic field. Figure 8a shows a single beam and figure 8b shows opposite beams.

DETAILED DESCRIPTION OF THE EMBODIMENTS

Introduction

The system presented in this paper is a 1.5 T MRI scanner integrated with a 6 MV radiotherapy accelerator. In essence, the design is a conventional 1.5 T MRI scanner with a small, single energy (6 MV) accelerator rotating around it. An outline impression of the integrated MRI accelerator is shown in figure 1. The system will allow simultaneous irradiation and MR imaging. MRI has several advantageous features for treatment guidance as well as for treatment monitoring.

1.1. Position verification Precise image guided radiotherapy has to deal with the uncertainties in the exact daily positioning of the radiation fields with respect to a changing anatoMV. The idea of integrating MRI functionality with an accelerator originates from the need for improved position verification by better visualising both tumour and surrounding structures. There are numerous competing position verification systems in clinical use or under investigation.

Megavoltage (MV) portal imaging is of great value for the introduction of conformal radiotherapy. A drawback of this method is that it visualises mainly the bony structures (Gilhuijs et at., 1996) which are not always correlated with the tumour position. Implanted fiducial markers correlate directly with tumour position and can be tracked automatically (Wu et at., 2001; Nederveen et at., 2001, 2003). This clearly improves the position verification for a limited number of tumour locations but this technique does not reveal the tumour shape or surrounding structures.

The image contrast can be improved by using additional kilovolt (kV) xray tubes, instead of using the MV treatment beam for imaging. This can be done 2-dimensionally (Drake et at., 2000) or 3-dimensionally by using multiple kV tubes and cameras (Shirato et at., 2000). More sophisticated is the integration of cone beam CT functionality with an accelerator (Jaifray et at., 2002; Berbeco et at., 2004). This provides great detail on the bony anatoMy or markers, but x-ray based soft-tissue contrast is inherently limited (Groh et at., 2002). Another position verification modality is the tomotherapy concept (Mackie et at., 1993) with megavoltage CT imaging integrated with an accelerator. This allows image reconstruction before treatment (Ruchala et at., 1999) but also has the advantage that the modulated beam can be used for image reconstruction during radiation treatment (Ruchala et at., 2000).

Ultrasound based position verification has the potential to visualise the tumour position, tumour shape and the surrounding structures. Despite some positive reports (Morr et at., 2002) the BAT ultrasound position verification technology is hampered by difficulties in evaluating the ultrasound images in relation with the pre-treatment CT images and by tissue deformations as caused by the abdominal pressure of the ultrasound probe (Van den Heuvel et at., 2002; Langen et at., 2003).

MR imaging can reveal both tumour position, tumour shape and the surrounding structures. It yields unequalled soft-tissue contrast and both 3D and 2D imaging with arbitrary slice orientation. Images can be made within 0. 3 s, facilitating tracking of intra-fraction motion of for instance lung tumours (Plathow et at., 2004). MRI can provide optimal image quality for on-line (intra- fraction) position verification needed to maintain geometric conformality during the entire course of radiotherapy also in case of significant tumour regression, for instance when treating cervical tumours (Van der Heide et al., 2004).

Determination of the patient specific position uncertainties allows the use of patient specific planning target volumes, introduced as adaptive radiotherapy (ART) by Yan et al. (2000). The availability of daily, online position verification data allaws the introduction of daily (re-) optimisation of the treatment plan (Wu et at., 2002). This procedure can currently be done off-line, for on-line re- optimisation fast on-line treatment planning is required. For Gamma Knife radiosurgery treatment planning this is already feasible (Wu et al., 2003) , for external beam planning, increasing the re-optimisation process speed is a topic of research (Wu et at., 2004).

1.2. Biological imaging Another feature of MRI is its potential to visualise tumour physiology using fMRI and MR spectroscopy (Stubbs, 1999). Examples are perfusion (Law et al., 2004), oxygenation (Zhong et al., 2003), hypoxia and acidity (Gillies et al., 2002) and even apoptosis (Hakumaki and Brindle, 2003). All these parameters may be used for measuring tumour activity and treatment response. The quantitative relation between tumour activity, dose sensitivity and image intensity is a topic of current basic and clinical research.

Dedicated for radiotherapy planning, the value of MR spectroscopy was shown by revealing the intra-prostate tumour location using the choline/citrate ratio (Zaider et at., 2000; DiBiase et al., 2002; Kurhanewicz et al., 2002). Also the staging and treatment planning of gliomas benefits from MR spectroscopy (Nelson et at., 2002; Law et al., 2003; Hsu et at., 2004).

This feature of MRI can be used for two, interconnecting purposes, i.e. treatment guidance and treatment monitoring. With images of tumour characteristics at hand it becomes possible to adapt the dose prescription to the local dose sensitivity. Instead of aiming for a geometrically conformed homogeneous dose, a biological conform dose can be given (Ling et al., 2000).

IMRT techniques as currently used in the clinic can deliver such 3D sculpted dose distributions that follow the heterogeneity of the tumour characteristics (Chao et al., 2001; Alber et al., 2003). The quantitative relation between measured tumour characteristics and dose sensitivity requires detailed studies. In order to reveal this relation daily registration of the dose deposition with the actual anatoMy as well as with the biological images is required. Integrated MRI functionality with the accelerator facilitates this daily registration. Such sophisticated treatment monitoring is a pre-requisite for the next step in adaptive radiotherapy: daily biological conformality.

A more exotic promise of MRI is the field of molecular imaging (Weissleder and Mahmood, 2001; Gillies, 2002). In vivo visualisation of tumour specific receptors and gene expression is already feasible in an experimental setting (Weissleder et al., 2000). This field needs further exploration but its potential is great (Coleman, 2003).

As mentioned before, the integration of MRI functionality with an accelerator allows on-line, high soft-tissue contrast imaging for position verification of both tumour and surrounding structures.

2. Physical Design The preliminary design consists of a small accelerator rotating around a modified, closed-bore 1.5 1 MRI system (Philips Intera TM) (Lagendijk et al., 2002; Raaymakers et al., 2004a), see figure 1. The technical feasibility study has focussed on three topics: Magnetic interference of the MRI system and the accelerator.

The permanent magnetic stray field would normally disturb the operation of the accelerator. On the other hand, magnetisation of the moving accelerator components would lead to modulation of the 1.5 T magnetic field with resultant image distortion and ghost artefacts as possible consequences. Active magnetic shielding therefore has to be applied to magnetically uncouple the MRI system and the accelerator. In section 2.2 the implications of this active shielding for the MRI magnet system will be discussed. To limit both cost and complexity, the concept is based on a cylindrical closed-bore main field magnet, with the accelerator irradiating the patient from outside the MRI system.

Beam characteristics after transmission through the MRI system.

The choice for a closed-bore MRI system with the accelerator irradiating the patient from outside the MRI system implies beam transmission through the MRI system. This introduces the potential for both scatter and absorption heterogeneity. The MRI magnet system and the gradient coil system are thus modified in order to minimise both scatter and absorption heterogeneity, as will be discussed in section 2.2. The remaining distortion of the beam characteristics will be discussed in section 2.3.

Dose deposition kernel in 1.5 T The photon beam travels unperturbed in the magnetic field. However the actual dose deposition, an avalanche of secondary electrons, initially released by a Compton interaction, is affected by the presence of a magnetic field, especially at tissue-air interfaces. In section 2.4 the impact of the magnetic field on the dose deposition kernel will be discussed.

2.1. Design of the integrated MRI accelerator The preliminary design of the MRI accelerator is basically an accelerator mounted on a ring around a slightly modified MRI system (see section 2.2). A ring mounted gantry is already successfully used in the Tomotherapy system (Mackie et al., 1993). For the MRI accelerator a customised ring has to be made with a larger diameter (approximately 1.3 m radius). To minimise the overall diameter of the system a compact radiation head design can be used. An option is a radiation head based on the six-bank multi leaf collimator (MLC) (Topoinjak et al., 2004).

2.2. Required modifications to the MRI system Main magnet The magnetic interference has the potential to hamper the integrated system in two ways: the operation of the accelerator could be distorted by the presence of the magnetic stray field. In particular, this affects the gun section with its low energy electrons before full acceleration. Further, accelerator components might become magnetised and thus induce a distortion on the main

magnetic field.

In order to limit image distortion, the main field homogeneity has to be of the order of few ppm for the entire field of view. Magnetic interference can be minimised by applying active magnetic shielding. Modification of the coil configuration of a standard Intera 1.5 T magnet has been investigated. By slightly modifying the number of turns of the shield coil of a standard actively shielded magnet design, it is possible to achieve a very low magnetic field in a toroidal volume in the midplane of the MRI magnet in which the accelerator can be placed. This modification leads to a slight increase of the far field of the magnet, but the field at larger distances is still sufficiently low to allow installation in the vicinity of conventional accelerators (typically at lOm distance).

Additionally, the coil configuration of the main magnet can be modified in such a way that there are no super-conducting coil windings in the transversal mid-plane of the MRI system. Also the passive shimming system can be placed outside the central area. This means that the equivalent of approximately 6 cm of homogeneous aluminium remains in the beams eye view, severely reducing both scatter induction and absorption heterogeneity, see also section 2.3. This modified magnet is not much different in construction, and does not significantly compromise the central field quality, i.e. the new design would yield the same imaging performance as a standard Philips system.

Gradient coil system The magnet system, i.e. the cryostat with the superconducting coils, can be adjusted to create a beam portal without compromising its performance. Next in the beam are the gradient coils. A conventional gradient coil poses a severe scatter and absorption problem. Several patterns of thick copper conductors with gaps of a few mm in between are located on top of each other. When irradiating through this, intensity contrasts up to 80% at a spatial resolution of mm can be expected. Therefore also radiation portals in the gradient coils are required, and the impact of a 20 cm gap in the gradient coils was investigated. For the gradient in the axial direction (Z-direction), the windings can easily be moved out of the central area. The coils for the transversal gradients (i.e. X- and y - direction) needed significant modification, but suitable transverse coil geometries without conductors covering the central 20 cm were indeed found. Figure 2

shows magnetic field plots for a possible design.

These preliminary investigations show the technical feasibility of placing the gradient coils outside the beam's eye view. The remaining construction material inside the beam's eye view is the equivalent of approximately 3 cm aluminium. For this preliminary design, the image quality in the transversal mid- plane will be unaffected and the field of view in the axial direction will be limited to approximately 25 cm, which is sufficient for this application.

As well as the above benefits in terms of homogeneity in the beam's eye view, other unexpected benefits were obtained. Specifically, placing the coils of the main magnet and/or the gradient coils on either side of the beam portal provides what are, in effect, annular shielding elements either side of the beam.

This means that radiation scattered from parts of the MRI system can be absorbed by the coils in their capacity as shields, and (therefore) that the amount of further shielding can be reduced or even eliminated.

RF coil Finally the beam has to travel through the RF coil. A conventional coil is reasonably transparent for radiation, it consists of cm-wide copper rods and perspex ribs of 4 mm and 3 cm thickness respectively. For a mono-energetic 2 MeV (which is approximately the mean energy of the 6 MV photon beam as used in section 2.4) photon beam these introduce absorption heterogeneities of approximately 22% at a spatial resolution of mm. Homogenizing the perspex carrier construction would limit the heterogeneity to the copper strips, at present a contrast of approximately 14%.

The thickness of the standard copper conductors is mainly motivated by mechanical considerations. Due to the skin-effect, only a thin surface layer of less than 30 microns actually carries the RF currents. A coil optimized for minimum radiation inhomogeneity can be made from 30 micron copper foil, placed on a non- metallic carrier. Thicknesses of 50 microns or less are particularly preferred, although 100 micron conductors may be suitable. This practically nulls the absorption contrast.

The skin effect is the tendency for high frequency signals to be transmitted in a layer at the surface of the conductor only. The depth is given by the equation: d= / 2 V,11000) where Po is the magnetic permeability of vacuum, a is the electrical conductivity and w is the frequency times 2ir. For the copper conductors used in the example, the values are: po: 1,25x106 a: 5.95x107 cv: 4.02x108, for an operating frequency of 64MHz This gives a value for d of 8.2 microns as opposed to a maximum advisable thickness (above) of 100 microns. Thus, it seems that by staying within 15 times the skin depth, good x-ray transmission qualities can be combined with good rf coil properties. Our preferred thickness of 30 microns thus corresponds to less than roughly four or five times the skin thickness, taking engineering tolerances into account.

Another modification of the RE coil is the introduction of photon beam radiation shielding outside the beam's eye view. This may help to minimise the out-of- field scatter at the patient level, as generated higher up along the beam (see also section 2.3).

2.3. Beam characteristics In section 2.2 the construction of beam portals in the main magnet and the gradient coil system is discussed. After these modifications, the beam still has to travel through various structures, as follows: Main magnet The aluminium cryostat housing (8 and 5 mm for the outer and inner cylinder respectively). Inside the cryostat, four thermal shields, of three aluminium together with the 1 cm stainless steel carrier of the super-conducting coils (represented as the black cylinder in figure 3(a)) are present. Additionally the lower part of the cryostat contains helium filling.

Gradient coils The carrier of the gradient coils, 4 cm homogeneous epoxy (represented as the thick light gray cylinder in figure 3(a)).

RFcoil By approximate engineering, the RE coil can be made into a 1 cm thick Perspex cylinder (represented as the innermost white cylinder in figure 3(a)) with a 30 micron thick copper foil on top of it.

The above mentioned layers together constitute the equivalent of approximately 10 cm of homogeneous aluminium. The scatter induction as a function of field size has been measured for a homogeneous slab of 10 cm aluminium, an internal wedge and simulated for the approximated MRI geometry. The results are shown in figure 3(b).

For the measurements cross profiles were taken at 5 cm depth with a Farmer NE2571 ionisation chamber (Nuclear Enterprises Limited, UK) in a water phantom for both a 5x5, lOxlO and 20x20 cm2 field. Measuring at 5 cm depth means that mainly the scattered photons are detected. This is representative for the MRI accelerator since scatter contamination by electrons will be virtually absent due to the recoiling of the scattered electrons as soon as these enter the bore of the MRI, see also section 2. 4. The water phantom was placed at 130 cm from the source and in case of the aluminium slab, the mass was placed on the tray holder (62 cm from the source) of a Elekta SLi 20 accelerator.

Geant4 Monte Carlo simulations (Agostinelli, 2003) were conducted to determine the scatter induction of the layered configuration, shown in figure 3(a). Convolution of a pencil beam dose distribution is not allowed since the MRI accelerator geometry is only symmetric along the central axis of the system.

Therefore we used a paper sheet approach, where unidirectional photons, sampled from a realistic energy spectrum as also used in Raaymakers et at.

(2004b), were fired from a random spot across a line perpendicular to this central axis. The length of this line determines the field width. The resulting paper sheet dose distribution was convoluted to the desired field size.

The scatter induction was determined at 10 cm from the field edge (i.e. the SO% iso-dose line). The scatter induction increases linearly with the field size. The scatter induction in the layered configuration is lower than for the case of a single slab and is comparable to the scatter induced when using an internal wedge. At this stage of the project, this seems an acceptable amount of scatter.

If necessary, the mass in the beams' eye view can be further reduced by defining a discrete number of gantry positions. Then for each gantry position the creation of a beam portal in both the carrier of the superconducting coils in the cryostat, the thermal shields and the carrier of the gradient coils can be created.

Such modifications to the MRI system are only allowed when they do not affect the mechanical rigidity of the system, so that the MRI performance will not be compromised. Another possible measure is including radiation shielding (out of the beam's eye view) in the RF coil, as mentioned above in section 2.2.

An explicit aim of the modification of the MRI components is to minimise the absorption heterogeneity. Any remaining heterogeneity will be dealt with using IMRT and/or compensators. For specific gantry angles a source of heterogeneity is the helium level in the cryostat. Transmission through a maximum of 15 cm liquid helium can occur, this will induce extreme ultraviolet fluoresence (McKinsey et aL, 2003) and more important for our purpose 10% intensity reduction. Another source of heterogeneity may be the omission of a flatness filter in order to minimise scatter induction and increase the accelerator output. Finally there is the minor absorption heterogeneity due to the divergence of the beam: rays at the edge of the field travel slightly longer through the various structures than rays in the center of the beam. This leads to heterogeneities of approximately O.5%.

Compensating the absorption heterogeneity will be done using both compensators and IMRT techniques. The latter requires a large field, high resolution MLC such as the six-bank MLC (Topolnjak et a!., 2004).

2.4. Dose deposition kernel The dose deposition of a 6 MV beam in the presence and absence of a 1.5T magnetic field was modelled using the GEANT4 Monte Carlo package (Agostinelli, 2003). Also the impact of 1.1T on a lxi cm2 field was measured. It is a safe approximation to neglect the impact of the magnetic gradient fields and the magnetic component of the RF field since these are at least 2 orders of magnitude lower than the main magnetic field. The impact of the 1.5 T transverse magnetic field on the dose deposition is described in detail by Raaymakers et al. (2004b). A pencil beam with a realistic energy spectrum enters a homogeneous half-space of water, with the magnetic field perpendicular to the beam axis, see figure 4. Figure 5 shows the cross-section for the pencil beam with the presence and absence of 1.5 T. Clearly the magnetic field induces an asymmetry in the dose deposition. By convolution the impact for larger fields can be calculated. In figure 6 dose profiles at 5 cm depth for a lxi and 5x5 cm2 field with the presence and absence of 1.5 T are shown. From figure 6(a) clearly the penumbra is affected, less clear is the shift of the entire field and the fact that the field width is unaffected. From figure 6(b) a decrease of the build-up distance can be seen when a magnetic field is present. In summary, for radiation fields where lateral electron equilibrium exists, the impact of the 1.5 T magnetic field on the dose deposition by a 6 MV photon beam is as follows: a decrease of the build- up distance by 5 mm, a symmetric increase of the penumbra in the direction perpendicular to the magnetic field by 1 mm and in the same direction a shift of the entire radiation field by approximately 0.7 mm. For fields too small to have lateral electron equilibrium in the central part of the field, the penumbra in the direction perpendicular to the magnetic field becomes asymmetric (See Raaymakers et al. (2004b) for details). These findings are mainly to determine whether the impact of the magnetic field on the dose deposition is endangering the technical feasibility of the MRI accelerator. The exact value of the results in case of tissue heterogeneities and a divergent beam might slightly differ fromthose obtained using a pencil beam approach in a homogeneous phantom.

The presence of a magnetic field does not affect the abutment of radiation fields. For treatment planning, the increased penumbra and reduced build-up distance can be incorporated in a conventional manner. The induced 0.7 mm shift of the isodose line in the direction perpendicular to the magnetic field requires the inclusion of asymmetric and direction-dependent dose profiles.

More dramatic is the impact at tissue-air interfaces due to the so called Electron Return Effect (ERE) (Raaijmakers et al., 2004). Electrons (as part of the electron avalanche) leaving the tissue and entering air, will be recoiled by the lateral magnetic field. With respect to scatter contamination (see section 2.3) the ERE is advantageous, since it will effectively purge the beam of scatter electrons. For patients, the ERE induces an increase of the dose at the distal side. Using multiple gantry angles as already done by conventional IMRT can solve the problem of overdosing the skin at the exit side of the beam.

The same affect occurs at air cavities. The ERE is shown for a water phantom with an air slit as schematically shown in figure 7. In figure 8(a) the dose profile on the central axis shows a clear local increase of the dose at the proximal side of the air cavity and at the distal side of the phantom. Raaijmakers et al. (2004) showed that for simple geometries such as this air slit, a reasonably homogeneous dose distribution can be achieved using opposite beam techniques, as shown by the dose profile in figure 8(b) . This is not a generic solution, for instance for a cylindrical air tube of 2 cm diameter (an idealised representation of a trachea) a box technique (with 4 beams) is required to homogenise the dose distribution inside the air cavity. For realistic air cavities IMRT techniques will be required. Currently we are investigating whether ERE can be exploited for IMRT treatments, i.e. for high dose delivery to small target volumes inside air cavities, as can be desired in head and neck tumours.

3. Technical and clinical challenges 3.1. Technical challenges In the summary of the physical design (section 2 above), various potentially insurmountable problems were addressed and we showed that integrating an MRI with an accelerator is technically feasible. The next step is coming to a conceptual design and then to a prototype. First the preliminary results need further and more detailed investigation. Monte Carlo simulations will be applied for quantifying the scatter induction in the presence of the magnetic field and in the actual MRI accelerator configuration. Also the impact of the altered dose deposition kernel for clinical IMRT in the presence of air cavities will be addressed. Additionally, various separate topics need to be investigated to fulfill the aim of the MRI accelerator, i.e. accurate radiation delivery at sub-mm resolution. These will be discussed below.

Absolute positioning When using MRI images for treatment guidance it is crucial that the images are geometrically correct and that the exact relation between the MRI coordinate system and the accelerator coordinate system is known. At our department a full quality assurance system for correction of MR images has been developed (Bakker et al., 1992, 1993, 1994; Bhagwandien et al., 1992, 1994; Moerland et al., 1995). System dependent image distortions could be reduced from 13 mm to 2 mm (Moerland, 1996) and patient induced distortions could be corrected to the order of the image pixel size (Bhagwandien et al., 1994; Moerland, 1996). These corrections and dedicated MR scanning procedures were developed explicitly for radiotherapy treatment planning purposes.

Geometrically correct MR images alone do not suffice for on-line MRI guidance. Also the precise relation between the coordinate system of the MRI and the accelerator needs to be determined. The spatial relation between the accelerator and the MRI system housing can be determined by off-line calibration. However the exact spatial relation of a MR image with the housing of the MRI system is dependent on the magnetic field distortions in the MRI system. These distortions are the sum of the imperfections of the MRI hardware, which can be compensated for (Moerland et al., 1995) and the distortions induced by heterogeneous susceptibility distribution (Lildeke et al., 1985). The latter are induced by the object to be scanned, i.e. these distortions are patient specific. Susceptibilty artifacts can be minimised using strong gradient fields (Bakker et al., 1992), or corrected for by using the calculated field distortions induced from the susceptibility distribution of a specific patient anatoMV (Bhagwandien et al., 1994; Burkhardt et al., 2003).

An elegant and exact correction can be done using the fact that the susceptibility artifacts occur in the direction of the read-out gradient (Chang and Fitzpatrick, 1992). This means that switching the read-out gradient causes the image distortion to be mirrored around the unperturbed position. By creating a distinguishable landmark in the image, for instance a marker in the table, and making two images with oppposite gradient directions, the unperturbed position - 18- of this landmark can be determined. This way the spatial relation between the MR images and the MRI system housing can be determined. As mentioned, this method yields exact results for a distinguishable landmark however the problem for an arbitrary image is that one can not define corresponding points in the two distorted images. (Chang and Fitzpatrick, 1992) looked for corresponding edges in the images and thus corrected the entire image. Basically what is needed is a deformable image registration between the two images, with the restriction that the deformation is only in the direction of the read-out gradient. The ITK toolkit (http://www.itk.org) can be used to investigate how non-rigid image registration can be applied in order to correct the entire image.

Dosimetry in 1.5 T The feasibility and calibration of dosimetry equipment in the presence of 1.5 1 has to be studied. Monte Carlo simulations will be used to quantify the behaviour of ionisation chambers in a magnetic field. The integration of MRI functionality with the accelerator urges the use of gel dosimetry for verification of IMRT dose distributions (Vergote et al., 2003; Gustavsson et al., 2003).

However, gel dosimeters based on radiation induced polymerization have an inherent slow (stable) response of 3 to 4 days (Mc3ury et al., 1999). On the other hand, gels infused with the Fricke solution have a rapid response, the ferrous (Fe2) ions are instantaneously converted to ferric (Fe3) ions, thereby altering the magnetic relaxation properties (Gore et al., 1984). The spatial resolution of these gels suffers from the diffusion of the ferric ions after irradiation (Schulz et al., 1990). Our MRI accelerator provides simultaneous irradiation and imaging, thus limiting the diffusion of ferric ions and the resulting deterioration of dose gradients.

3.2. Clinical challenges The MRI accelerator combination can be used both for position verification and for measuring treatment response. It is expected that the first clinical benefits will flow from improved position verification (on-line and soft-tissue visualisation).

A different challenge for position verification is when not only the position but also the shape of the clinical target volume is subject to daily changes, for instance for cervical radiotherapy. For the latter, the clinical target volume (CIV) consists of the primary, regressing tumour and a region suspected of microscopic invasion. Within the CIV some parts (e.g. the uterus) are expected to move substantially, while other parts (e.g. the lymph nodes) are expected to be relatively fixed to the bony anatoMy. Moreover, tumour regression may cause a substantial shrinkage of part of the CIV during the course of the treatment. As a result the uncertainties for cervix radiotherapy are an order of magnitude larger than for prostate radiotherapy. On-line MRI guidance offers the potential to realise a highly conformal plan, and to maintain this conformality throughout the course of radiotherapy.

Yet another opportunity is the tracking of intra-fraction motion of (for instance) lung and liver tumours. Present studies investigate tracking of the breathing cycle with cone beam CT (Zijp et al., 2004). On-line MRI facilitates the tracking of breathing related motion, including its irregularities (Plathow et al., 2004).

These position verification applications and the increasing use of anatoMV MR images for radiotherapy treatment planning will facilitate the gradual introduction of more complex, functional and biological MRI data. First as an aid for treatment planning, and if proven useful, later for adaptive treatment guidance or measuring treatment response, with on-line adaptive radiotherapy using the MRI accelerator as the ultimate aim.

4. Conclusion

Preliminary investigations show the feasibility of an integrated MRI accelerator system. Such a system will facilitate daily position verification and treatment adjustment. Ultimately the system may be used for daily treatment optimisation by providing the imaging information for on-line treatment planning. On-line MRI may also provide treatment monitoring and treatment response assessment required for further biological optimisation. By doing so, this system can facilitate true MRI guided radiotherapy (MRIgRT).

It will of course be understood that many variations may be made to the above-described embodiment without departing from the scope of the present invention.

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Claims (12)

1. Radiotherapeutic apparatus comprising a linear accelerator that is rotateabje around an isocentre, and a magnetic resonance imaging apparatus disposed around the isocentre, the rf coil windings of the magnetic resonance imaging system being of a conductive material shaped as a strip of thickness not greater than dmax where: d=15/ 2 V /ioaa) where Po is the magnetic permeability of vacuum, a is the electrical conductivity and w is the frequency times 2ir.

2. Radiotherapeutic apparatus according to claim 1 in which the conductive material is copper.

3. Radiotherapeutic apparatus according to claim 1 or claim 2 in which the rf coil windings have a thickness less than 100 microns.

4. Radiotherapeutic apparatus according to claim 1 or claim 2 in which the rf coil windings have a thickness less than 50 microns.

5. Radiotherapeutic apparatus according to any one of the preceding claims in which the thickness of the rf coil windings is less than: d=5f 2 I,LIoaO)

6. Radiotherapeutic apparatus according to any one of the preceding claims in which the rf coil windings are mounted on a non-metallic carrier.

7. Radiotherapeutic apparatus according to claim 6 in which the nonmetallic carrier is PerspexrM.

8. Radiotherapeutic apparatus comprising a linear accelerator that is rotateable around an isocentre within a beam transmission plane, and a magnetic resonance imaging apparatus disposed around the isocentre, the magnetic resonance imaging apparatus including coils for generating a magnetic field, disposed in an annular ring located in a plane displaced from the beam transmission plane.

9. Radiotherapeutic apparatus according to claim 8 in which coils of the magnetic resonance imaging apparatus are placed in a first annular ring displaced from the beam transmission plane in a first direction and a second annular ring displaced from the beam transmission plane in a second direction opposite to the first direction.

10. Radiotherapeutic apparatus according to claim 9 in which the coils of the magnetic resonance imaging apparatus disposed in the first and second annular ring include coils of the gradient coil system.

11. Radiotherapeutic apparatus according to claim 9 or claim 10 in which the coils of the magnetic resonance imaging apparatus disposed in the first and second annular ring include coils of the main magnet system.

12. Radiotherapeutic apparatus substantially as herein disclosed with reference to and/or as illustrated in the accompanying figures.