WO2014133391A1

WO2014133391A1 - Dipole antenna for a magnetic resonance imaging system

Info

Publication number: WO2014133391A1
Application number: PCT/NL2014/050125
Authority: WO
Inventors: Alexander Jan Eberhard Raaijmakers RAAIJMAKERS; Cornelis Antonius Theodorus Van Den Berg; Dennis Wilhelmus Johannes KLOMP; Peter Rudolf LUIJTEN
Original assignee: Umc Utrecht Holding B.V.
Priority date: 2013-03-01
Filing date: 2014-03-03
Publication date: 2014-09-04
Also published as: EP2962361A1

Abstract

The invention relates to an antenna array for transmitting and receiving RF radiation to and from a target at a predetermined depth in an object. The antenna array can be used in a magnetic resonance image system. The antenna array is provided with at least one dipole antenna (401) comprising two legs, and each leg comprises at least two segments (401, 420; 404, 405) and a coupling device (406, 407) arranged to electrically connect the respective segments.

Description

Dipole antenna for a Magnetic Resonance Imaging System

TECHNICAL FIELD

The invention relates to an antenna array in a magnetic resonance imaging system.

A conventional MRI scanner typically comprises a coil system arranged to generate the static magnetic field B₀. Next to that, the gradient coils enables spatial encoding on the B₀ field which is required for tomographic imaging. Furthermore, the MRI system comprises an RF transmit and receive system to generate and pick-up the Bi fields with a first frequency. These Bi fields may excite the atomic nuclei of an object within the coil system. Upon relaxation, the nuclei emit signals of the first frequency that are being picked-up by the RF transmit and receive system.

The frequency of these Bi fields is proportional to the B₀ field strength, e.g. 64 MHz at a Bo field of 1.5 T for excitation of hydrogen nuclei.

Over the past decades, MRI systems have developed by an ever increasing magnetic field strength of the B₀ field. A higher field strength of the B₀ field provides more signal and, therefore, better image quality.

Since 2004, human whole body MRI systems at 7 Tesla exist and recently 9.4 T and 11.7 T systems have been installed. Conventional MRI systems in hospitals, however, currently operate at a B₀ field strength 1.5 or 3 Tesla. These conventional

MRI systems comprise a so-called body coil for generating the Bi-field that excites the atomic nuclei of the object in the MRI system. The body coil is integrated within the bore of the MRI system. Presently, the body coil can generate a homogeneous Bi field all over the imaging region for the MRI system operating at a B₀ field of 1. 5 T.

However, body coils for use in MRI systems operating at a B₀ field of 3 Tesla, due to the higher frequency of the Bi field, show substantial inhomogeneities in the Bi-field. Body coils for use in MRI system operating at a B₀ field of 7 Tesla or higher, are no longer feasible because the inhomogeneities in the Bi field are so severe that they lead to unwanted effects, for example signal voids in constructed images of the receive signals.

Instead of these body coils, MRI systems operating at higher B₀ fields, for example, 7 Tesla, comprise local transmit coils for transmitting the Bi field. The local transmit coil is a separate coil that is not integrated within the MRI system like a conventional body coil. Instead, the local transmit coil can be placed on the scanner table and/or on the target to be imaged in the MRI system.

If the target is relatively small, such as the human head, then the local transmit coil may have a similar design of the body coil, but just smaller. This design is known as 'birdcage coil'. Presently, imaging of the human head using MRI systems operating at a Bo field of 7 Tesla is performed routinely with a so-called birdcage coil as an example of a local transmit coil.

Imaging of the human head using MRI systems operating at a B₀ field of 7 Tesla has demonstrated superior image quality in comparison to imaging of the human head using MRI systems operating at B₀ fields of lower field strengths.

However, imaging of other targets within the body, for example, a heart or a prostate, requires a different design because the birdcage design would become too big and essentially behaves like a body coil with inherently, a low efficiency and signal voids due to interference of the Bi fields. Instead, the special designed transmit coils are combined in an array comprising multiple transmit elements, a so-called transmit array. These elements are distributed around the body at the desired longitudinal position to achieve sufficient signal strength at the target to be imaged. The transmit elements can also be used as receive elements within the same MRI experiment, by which the array is called a transmit-receive or transceive array.

The MRI system comprising this transmit array is arranged to provide an RF signal to each element. The RF signal to the respective elements generates a Bi field with the required frequency and a predetermined phase and amplitude to each element. In this way, the signal interference of the respective Bi fields can be modulated to ensure constructive interference of the fields in the region of interest.

These arrays can also be used in 3 Tesla systems to alleviate field

inhomogeneities .

However, some problems may occur in designing these array elements. Known requirements for these array elements are: a high efficiency in transmit mode, or equivalently in receive mode, high sensitivity for RF signals originating at a depth of the imaging target. In addition, the element design needs to take the safety issue of tissue heating into account. In general, any RF exposure to tissue results in energy absorption and consequently tissue heating. The distribution of energy absorption is known as the local specific absorption rate, SAR, distribution denoted in W/kg. Such a distribution, like the Bi field distribution, is a characteristic of an array element. Local peak values in the SAR distribution may give rise to excessive temperature rise. To take into account the alleviating effect of local heat conduction, the lg averaged or lOg averaged SAR distribution is evaluated. The maximum value within this distribution limits the duty cycle at which an MRI scan can be performed.

Therefore, the array elements should have a low maximum local SAR value. Finally, the elements should experience little mutual coupling, because this mutual coupling causes loss of power due to scattering losses.

EP2030720 discloses a substrate with a first side and a second side, wherein the substrate comprises a dielectric material, at least one dipole antenna, wherein the dipole antenna is attached to the second side of the substrate, wherein the dipole antenna comprises a first connection adapted for connecting the dipole antenna to a

transmission line.

A conventional dipole antenna comprises two legs of a conducting material and the antenna is fed with a balanced signal in the center of the legs, wherein the full length of the both legs is equal to half the wavelength at the operating frequency.

However, the conventional dipole antennas will generate high E-fields close to the antenna. This effect makes the conventional dipole antenna's less suitable for use close to objects because the high E-fields lead to high levels of energy absorption and, consequently, temperature rise of the skin and subcutaneous tissue.

A drawback of the element array can be an insufficient signal at a depth of the imaging target, in particular, with an acceptable SAR, that may deteriorate the image quality of the reconstructed images.

A further drawback may be that the sensitivity at depth is low. These drawbacks may also be present in MRI systems at lower field strength than 7 Tesla.

SUMMARY OF THE INVENTION

It is an object of the invention to improve an antenna for an MRI system.

According to a first aspect of this invention this object is achieved by an antenna array provided with at least one dipole antenna comprising two legs, and each leg comprises at least two segments and a coupling device arranged to electrically connect the respective segments. The invention is based on the insight that in MRI systems, an improved performance can be obtained by manipulating the currents and voltages on the antenna, for example, by shortening the length of the dipole with respect to half of the wavelength of the obtained RF field. By manipulating the length the voltage/ current ratio can be adapted such that the E/H ratio, or the local wave impedance matches the equilibrium impedance in, for example, the tissue of the target in a body to be imaged.

The shortening of the antenna enables tailoring of the field of view of the antenna array to the object being scanned. If the object is superficial and small, a short antenna is preferred. If the object is large and/or deep, a long antenna is preferred.

In MRI systems a B₀ field corresponds to a Larmor frequency of nuclei in the object to be imaged. The Larmor frequency is a characteristic frequency of a spin of a nuclei in an external magnetic field given by ω = γ B , wherein

ω represents the Larmor frequency,

γ represents the gyromagnetic ratio of the nuclei,

B is the strength of the magnetic field.

The nuclei can be, for example, protons, the corresponding Larmor frequency at a B₀ field strength of 1.5 MHz is 64 MHz. The RF transmitters and receivers of the MRI system are tuned to this Larmor frequency. Other nuclei are for example sodium or phosphorous. Presently, RF antennas or applicators in MRI systems using Larmor frequencies corresponding to higher B₀ field strengths reach a new regime wherein the conventional coil design loses its efficiency and/or causes high SAR level in the imaging object. It is well known that at high frequencies coils start to behave as antennas. Close to the coil, a region exists where the Bi field is very high. These are the reactive fields that are associated to the conventional desired resonance operation of the coil. But at larger depths inside the tissue, the reactive fields will decay and the, so- called, radiative fields become more dominant. Here, the RF field distribution has turned into an electromagnetic wave that propagates away from the coil. At high B₀ field strengths, generating a propagating electromagnetic wave is therefore a more suitable way of achieving the required strength of Bi field in the targets that are more deeply located in the body.

The principle of reciprocity holds in the design of receive and transmit antennas. So, what is stated above for transmit antennas also applies for receive antennas. The sensitivity pattern of a receive antenna is similar to the Bi field pattern of a transmit antenna. Preferably, the array elements for body imaging at 7 Tesla should not generate a large resonant field, because the shorter wavelength corresponding to the frequency of the Bi field results in a penetration depth to only a small region around the array element wherein the resonant field exist. The dipole antenna should aim to generate a far-field, i.e. to emit a propagating electromagnetic wave towards the imaging target. In general, an electromagnetic wave consists of both electric fields (E-fields) and magnetic fields (H-fields). In order to generate such a propagating wave, an antenna should provide both. This effect distinguishes an antenna from a coil. Conventionally, coils for use in MRI systems should be designed to avoid direct generation of E-fields since E-fields are responsible for SAR. At the frequencies used in MRI system operating at B₀ fields with higher field strength, however, SAR cannot be avoided since E-fields are locally induced by the time varying magnetic fields, resulting in enhanced SAR levels anyhow. On the contrary, a highly resonant coil may not generate E-fields directly, but the E-fields that are induced by the time varying magnetic fields result in much higher SAR. The above analysis results in a new approach for coil design. Note that this new approach can also realize increased efficiency/sensitivity for lower field strengths, for example, at 3 Tesla.

The antenna array comprises, for example, eight dipole antennas. The legs of the dipole antenna can be aligned along a line and arranged in opposite directions.

In an embodiment the segments can be evenly distributed along the length of the legs. Herewith, a symmetrical design of the dipole antenna is obtained. The coupling device can be applied to electrically match the impedance of the dipole antenna of a determined length with the impedance of an output of an RF amplifier.

In a further embodiment the coupling device comprises a switch for connecting the respective segments. By connecting the respective segments via the switches, the length of the legs can be dynamically adapted to a predetermined length, for example, the length of a longitudinal dimension of the target to be imaged or a predetermined length at which the antenna provide maximum signal/sensitivity with respect to the depth of the target in the body. The first switch may comprise a PIN diode or a FET. The switching of the coupling devices allow the antennas within the array to switch from a first mode to a second mode and vice versa. In a further embodiment the coupling device comprises an electrical network between the respective segments. By adapting the impedance of the electrical network an inductive or a capacitive characteristic of the coupling device can be obtained.

In a further embodiment the electrical network comprises an inductance and/or a capacitance. By using an inductive element or a capacitor a predetermined impedance of the coupling device can be obtained.

In a further embodiment the electrical network comprises, the switch, the switch being arranged to switch the network in a first mode or second mode, wherein, in a first mode, the electrical network has a first impedance and, in a second mode, the electrical network has a second impedance and the first impedance is different from the second impedance.

In a further embodiment the first impedance is a first inductive impedance and the second impedance is a second inductive impedance, and the first inductive impedance is lower than the second inductive impedance. This arrangement enables alternating between different current/voltage patterns in the dipole so that local maxima in the current/voltage pattern can be at different locations in the in dipole in the first and second mode.

In a different embodiment the first impedance is an inductive impedance and the second impedance is a capacitive impedance. For example, in the first mode the segments of the antennas can be coupled via inductive elements and in the second mode the segments of the antennas can be coupled via capacitors. In this example the first mode is characterized by a low SAR level and low sensitivity/efficiency, although the B1+ ratio/sqrt (SAR) is more beneficial and the second mode is characterized by a high SAR levels and high sensitivity/efficiency, particularly for shallow targets.

For example, when the antenna array is used in an MRI system, the first mode can be, for example, a transmit mode, for transmitting the Bl field to the target and the second mode can be a receive mode for receiving an RF field from the target. In the first mode a low SAR can now be obtained and in the second mode a high sensitivity for receiving the RF field from the target.

In a further embodiment the antenna array comprises a substrate arranged to support the legs. The substrate can be a printed circuit board, PCB. The substrate can also be adapted to a predetermined shape to match for example a part of a body to be imaged. In a further embodiment the inductive element comprises a strip of conducting material provided in a meander shape on the substrate. The conducting material can be, for example, copper. In other embodiments the strip of conducting material may have different shapes.

In a further embodiment the dipole antenna is integrated with the printed circuit board.

In a further embodiment the antenna array is further adapted to transmit electromagnetic waves for performing radio frequency hyperthermia treatment of the object. In this arrangement the antenna array can be used for magnetic resonance imaging and the same antenna array can be used for hyperthermia treatment. The temperature of tissue can now directly be measured using magnetic resonance imaging and the hyperthermia can be accurately guided and controlled.

The invention also relates to a magnetic resonance imaging system comprising a first magnet for generating a first magnetic field for orienting magnetic spins of nuclei, a magnetic field gradient coil for spatial encoding and manipulating the orientation of the magnetic spins of the nuclei, a radio frequency system for generating and receiving RF signals; an antenna array as claimed in any of the claims 1- 13, connected to the RF system, and a controller arranged to control the magnetic field gradient coil and the radio frequency system to collect magnetic resonance image data and to reconstruct images from the collected magnetic resonance image data, wherein the controller is further arranged to switch, in a first mode, a coupling device between two adjacent segments in the dipole antenna to a first impedance, and in a second mode, to a second impedance, which second impedance is different from the first impedance. Brief description of drawing

The above and other, more detailed aspects of the invention will be elucidated and described hereinafter, by way of example, with reference to the accompanying drawing. Therein:

Fig. 1 shows a first embodiment of the antenna array;

Fig. 2 shows a first embodiment of an element of the antenna array;

Fig. 3 shows a second embodiment of an element of the antenna array; Figs. 4A, 4B show different examples of connection circuits between transmission lines and the antenna array;

Figs. 5 A, 5B, 5C show different embodiments of a coupling device;

Fig. 6 shows diagrammatically an embodiment of an MRI system; and

Fig. 7A, 7B, 7C, 7Dshows diagrams of simulations of an expected SAR and Bl field as a function of depth in an object.

DETAILED DESCRIPTION OF THE PRESENT INVENTION Figure 1 shows an MRI system 100 comprising an RF system 101 provided with an antenna array comprising a plurality, for example, eight antenna elements 110. The MRI system 100 further comprises an RF system 101, RF amplifiers 102 and transmit/ receive switches 103 and a control system 130. Each antenna element 110 comprises a dipole antenna. Each antenna element 110 has its own transmission line 115 connecting it to a RF system 101. The RF system 101 is arranged to independently control the amplitude and phase of the RF signals applied to the antenna elements 110. The control system 130 controls the RF generator 101.

Fig. 2 shows a first embodiment 200 of an antenna element that can be used in the MRI system. Antenna element 200 comprises a dipole antenna. In this example, the length is 30 cm. Furthermore, the dipole antenna comprises two legs 220,221, wherein each leg is divided in two segments 201,202;203,204. The segments can be equally distributed along the legs. Each leg comprises a conductive strip, the conductive strips of both legs are aligned along an axis.

The segments 201,202,203,204 can be mounted on a substrate 230, for example, a printed circuit board, PCB. The thickness of the PCB is, for example, 2 mm.

Furthermore, the segments 201,202; 203,204 of each leg can be coupled via a coupling device 205, 206. The coupling device may comprise an electrical network with a predetermined impedance. For example, the electrical network may comprise an inductive element consisting of, for example, a conductor consisting of a meander shaped conductive element. The predetermined impedance of the coupling device can be matched to the output of the RF amplifiers. Furthermore, connections 207,208 are provided at each segment, to receive an RF signal of a matched and balanced power source and/or to provide signals to the receive system Fig. 4 A shows an embodiment of a connection circuit between a transmission line 400 and a dipole antenna 401 of the antenna array. The segments 402,403; 404,405 of the dipole are connected to each other via an inductance 406,407 . The connection between the transmission line and the segments of each leg of the dipole antenna can be made via a first circuit comprising a first capacitance 408, and a second capacitance 409, wherein the first capacitance 408 is connected in series with a central conductor of the transmission line, the second capacitance 409 is connected in series with an outer conductor of the transmission line 400.

Fig. 3 shows a second embodiment 300 of an antenna element that can be used in an MRI system. In this example the legs 320,321 are divided in three segments

301,302,303; 304,305,306 of 5 cm length each. The total length of the segment is then 30 cm. Each segment may comprise a conductive strip, the conductive strips of both legs are aligned along an axis. The segments are equally distributed along each leg on the substrate 330, for example a printed circuit board, PCB. The segments

301,302,303;304,305,306 can be coupled via a coupling device 307,308; 309,310. The coupling device may comprise, for example, an inductive element comprising for example, a meander shaped conductive element, and/ or a capacitor. Each of the legs of the dipole antenna can receive an RF signal via the connections 311, 312. An RF system can be coupled to the antenna array via a transmission line connected to the dipole antenna.

Fig. 4B shows a second embodiment of a connection between a transmission line and a dipole antenna of the antenna array. The segments 422,423,424; 425,426,427 of the dipole are connected to each other via capacitors 428,429,430,431. The connection between the transmission line 400 and each leg of the dipole antenna is made via a first circuit comprising a first inductor 432, a second inductor 433 and a capacitor 434, wherein the first inductor 432 is connected in series with a first central inductor of the transmission line 400, the second inductor 433 is connected in series with an outer conductor of the transmission line 400 and a capacitor 434 is connected between the first and second connections 207,208 of the dipole antenna .

Fig 5 A shows a first embodiment of a coupling device 500 comprising an electrical network of a first branch 501 comprising an inductance 502 and a second branch 503 comprising a capacitance 504, wherein the respective branches 501,503 are parallel connected in the electrical network. Fig. 5 B shows a second embodiment of a coupling device 510 that can be used in coupling of the segment 201,202 comprising an electrical network of a first branch 511 comprising an inductance 512 and a switch 515, for example, a PIN diode and a second branch 510 comprising a capacitance 514 The PIN diode can be switched by a bias voltage to change the impedance of the electrical network 510 from high to low or vice versa in order to connect or disconnect the respective segments in the legs of the dipole antenna 201, 202 . Alternatively, the elements in the electrical network can be chosen such that the network switches from capacitive to inductive and vice versa.

Fig. 5 C shows a third embodiment of a coupling device 520 that can be used to switch the antenna between a first and a second mode.. The coupling device 520 comprises an electrical network of a first branch 521 comprising a first inductance 522 , and a second electrical network, and a second branch 523 comprises a first capacitance 524. The second electrical network comprises, in a first branch, a second inductance 527 and, in a second branch 528, a second capacitance and a switch 525 connected in series with the capacitance 529 . The switch, for example, a PIN diode or FET can be switched by a bias voltage to change the impedance of the first network from capacitive to inductive or vice versa or to connect or disconnect the respective segments 201,202 of the legs of the dipole.

In this embodiment the different modes of the coupling devices

307,308,309,310 can be applied to select the length of the legs 320,321 by adding or isolating the segments 301,302;305,306. In this way the length of the antenna element can be adapted to the length of an envisaged target of the body of the patient or a depth of an envisaged target in the body of the patient.

The antenna array can also be adapted for transmitting electromagnetic waves at different frequencies.

In an embodiment the respective segments 301,302; 303,304; 305,306 can be adapted to dipole antennas that are connected to respective RF receivers.

Furthermore, in a transmit mode the total length of the legs of the antenna element can have a different length than the length of the legs of the antenna element in a receive mode by connecting one or more segment in each legs.

In an embodiment two modes can be possible. A first mode wherein all the adjacent segments in the respective legs are connected to each other, and a second mode wherein one or more adjacent segments in each leg are disconnected from a first segment in the leg that is connected to the transmission line.

In an embodiment the substrates 201, 315 can be provided with a spacer of PMMA to maintain a specified distance between the antenna elements and the body. The thickness can be for example 8 mm.

Fig. 6 shows an embodiment of a magnetic resonance imaging system. An object to be imaged, for example a body of a patient 611, is within the bore of a cylindrical coil 606 for generating a magnetic field for orienting magnetic spins of nuclei, for example protons. The magnet generates a magnetic field sufficient for acquiring a magnetic resonance imaging image within an imaging zone. For example, 3 Tesla or 7 Tesla. The magnet 606 may comprise a permanent, electromagnet or a superconducting magnet. The magnet 606 can be a cylindrical magnet with a bore adapted for receiving the patient, or it can be a split magnet design with two cylindrical magnets located coaxially next to each other. The magnet can also be a split coil or so a called open magnet design. Also, inside of the bore of the magnet are the magnet field gradient coils 608 for spatial encoding and manipulating the orientation of the magnetic spins of the nuclei, the magnetic field gradient coils 608 are connected to the magnet field gradient power supply 604.

An antenna array according to an embodiment of the invention is arranged adjacent to the patient 611 In this example two dipole antennas 609, 610 are shown. Transmission lines 612 connect each dipole antenna 609, 610 to a radio frequency transmit/receive system 605. The radio frequency system 605 and the magnetic field gradient power supply 604 are connected to a controller 601. The controller 601 comprises, for example a microprocessor which is adapted for executing a computer program product stored in a memory 603. In this example the computer program product is arranged to control the magnetic field gradient coil 608 and the radio frequency transmit/receive system 605 to collect magnetic image resonance data and to reconstruct images from the collected magnetic resonance image data. The controller is further arranged to control the coupling devices, in a first mode, to switch the segments to a first impedance , and in a second mode, to switch the segments to a second impedance. For example, by controlling the bias currents to each of the PIN diodes in the coupling devices. The controller 601 also comprises a user interface 602 which can be used by an operator to control the magnetic resonance imaging system. The first mode can be a transmit mode and the second mode can be a receive mode of the antenna array. In another embodiment the first and second mode can be both transmit modes.

Fig. 7 A shows simulation results of a30 cm antenna comprising 2 legs of each 3 segments as described above. In the diagram the lines 701,702, 703, 704, 705, 706 correspond to a calculated Bl field as a function of depth in a phantom body at different arrangements of the coupling device. Graphs 701, 702,703 correspond to a coupling device comprising a coil of respectively a value of 150 nH, 100 nH, 50 nH, Graph 704 corresponds to a conventional dipole and graphs 705,706 correspond to a coupling device comprising a capacitor of respectively a value of 5pF and 1 pF.

Fig. 7B shows simulation results of a 30 cm antenna comprising 2 legs of each 3 segments as described above, in the diagram the lines 711,712, 713, 714, 715, 716 correspond to a calculated SAR level along the longitudinal axis of the dipole at the surface of a phantom body for different arrangements of the coupling device. Graphs 711, 712,713 correspond to a coupling device comprising a coil of respectively a value of 150 nH, 100 nH, 50 nH, Graph 714 corresponds to a conventional dipole and graphs 715,716 correspond to a coupling device comprising a capacitor of respectively a value of 5pF and 1 pF.

Fig. 7C shows simulation results of a 30 cm antenna comprising 2 legs of each 3 segments as described above, in the diagram the lines 721, 722, 723, 724, 725, 726 correspond to a calculated Bi field relative to a conventional dipole, as a function of depth in a phantom for different arrangements of the coupling device. Graphs 721, 722, 723 correspond to a coupling device comprising a coil of respectively a value of 150 nH, 100 nH, 50 nH, Graph 724 corresponds to a conventional dipole and graphs 725, 726 correspond to a coupling device comprising a capacitor of respectively a value of 5pF and 1 pF.

Fig. 7 D show simulation results of a 30 cm antenna comprising 2 legs of each 3 segments as described above, in the diagram the lines 731, 732, 733, 734, 735, 736 correspond to the calculated ratio between to Bl and the square root of SARma_X as a function of depth in a phantom body for different arrangements of the coupling device. All graphs are plotted relative to a conventional dipole. Graphs 731, 732, 733 correspond to a coupling device comprising a coil of respectively a value of 150 nH, 100 nH, 50 nH. Graph 734 corresponds to a conventional dipole and graphs 735, 736 correspond to a coupling device comprising a capacitor of respectively a value of 5pF and 1 pF.

It should be noted that the above-mentioned embodiments illustrate rather than limit the invention, and that those skilled in the art will be able to design many alternative embodiments without departing from the scope of the appended claims. Other variations to the disclosed embodiments can be understood and effected by those skilled in the art in practicing the claimed invention, from a study of the drawings, the disclosure and the appended claims, In the claims the word "comprising" does not exclude other elements or steps, and the indefinite article "a" or "an" does not exclude a plurality. A single processor or other unit may fulfil the functions of several item recited in the claims. The mere fact that certain messages are recited in mutually different dependent claims does not indicate that a combination of these measures cannot be used to advantage. A computer program may be stored/distributed on a suitable medium, such as an optical storage medium or a solid-state medium supplied together with or as part of other hardware, but may also be distributed in other forms, such as via the Internet or other wired or wireless telecommunication systems. Any reference signs in the claims should not be construed as limiting the scope.

Claims

1. Antenna array for transmitting and receiving RF radiation to and from a target, wherein the antenna array is provided with at least one dipole antenna comprising two legs, and each leg comprises at least two segments and a coupling device arranged to electrically connect the respective segments.

2. Antenna array as claimed in claim 1, wherein the at least two segments are equally distributed along the leg.

3. Antenna array as claimed in claim 1 or 2, wherein the coupling device comprises a switch for connecting the respective segments.

4. An antenna array as claimed in claim 1 - 3 , wherein the coupling device comprises an electrical network between the respective segments.

5. An antenna array as claimed in 4, wherein the electrical network comprises an inductance and/or a capacitance.

6. Antenna array as claimed in claim 3 - 5, wherein the electrical network comprises the switch, the switch being arranged to switch the network in a first mode or second mode, wherein, in a first mode, the electrical network has a first impedance and, in a second mode, the electrical network has a second impedance and the first impedance is different from the second impedance.

7. Antenna array as claimed in claim 6, wherein the first impedance is a first inductive impedance and the second impedance is a second inductive impedance, and the first inductance impedance is lower than the second inductive impedance.

8. Antenna array as claimed in claim 6, wherein the first impedance is an inductive impedance and the second impedance is a capacitive impedance.

9. Antenna array as claimed in any of the claims 1- 8, wherein the antenna array comprises a substrate arranged to support the legs.

10. Antenna array as claimed in any of the claims 9, wherein the inductive element comprises a strip of conducting material provided in a meander shape on the substrate.

11. Antenna array as claimed in claim 9 or 10, wherein the substrate is a printed circuit board.

12. Antenna array as claimed in claim 11, wherein the at least one dipole antenna is integrated with the printed circuit board.

13. Antenna array as claimed in any of the claims 3 - 12, wherein the first switch comprises a PIN diode or FET.

14. Antenna array according to any of the claims 1 - 13, wherein the antenna array is further adapted to transmit electromagnetic waves for performing radio frequency hyperthermia treatment of the object.

15. Magnetic resonance imaging system comprising

a first magnet for generating a first magnetic field for orienting magnetic spins of nuclei;

a magnetic field gradient coil for spatial encoding and manipulating the orientation of the magnetic spins of the nuclei;

a radio frequency, RF, system for generating and receiving RF signals;

an antenna array as claimed in any of the claims 1- 13, connected to the RF system; and

a controller arranged to control the magnetic field gradient coil and the radio frequency system to collect magnetic resonance image data and to reconstruct images from the collected magnetic resonance image data, wherein

the controller is further arranged to switch, in a first mode, a coupling device between two adjacent segments in the dipole antenna to a first impedance , and in a second mode, to a second impedance, which second impedance is different from the first impedance.