WO2011135312A2

WO2011135312A2 - Mri rf coil with improved pin diode switch and reduced b1 distortions

Info

Publication number: WO2011135312A2
Application number: PCT/GB2011/000667
Authority: WO
Inventors: Nandita Maria Desouza; David John Gilderdale; Maria Angelica Schmidt
Original assignee: Institute of Cancer Research Royal Cancer Hospital; Royal Marsden NHS Foundation Trust
Current assignee: Institute of Cancer Research Royal Cancer Hospital; Royal Marsden NHS Foundation Trust
Priority date: 2010-04-29
Filing date: 2011-04-28
Publication date: 2011-11-03
Anticipated expiration: 2012-10-29
Also published as: WO2011135312A3; GB201007198D0

Abstract

An MRI detector for detecting RF signals has a solenoid coil which completes one or more solenoid turns. The detector further has one or more tuning capacitors positioned in the coil for tuning the resonance frequency of the coil. The detector further has a PIN diode positioned in the coil and arranged in series with the tuning capacitors. The detector further has circuitry for controlling the bias of the PIN diode. When the PIN diode is forward biased the coil is able to detect RF signals emitted by spins from a subject in response to an excitation Bi magnetic field. When the PIN diode is zero or reverse biased the coil is substantially prevented from distorting the excitation magnetic field.

Description

MRI DETECTOR

The present invention relates to a magnetic resonance imaging (MRI) detector for detecting radio frequency (RF) signals.

The closer a magnetic resonance (MR) signal detector (receiver coil) is placed to its target, the better the signal-to-noise ratio (SNR) of the resultant data. SNR is maximized if the coil size is optimised for the region of interest, and coils located very close to the target to be studied can achieve dramatic improvements in SNR relative to conventional external enveloping coils. Internal access for such small local coils may be gained via natural orifices or minimally invasive techniques. D.J. Gilderdale et al., The British Journal of Radiology, 72 (1999), 1141-1151, reviews the design,

development and clinical applications of solid, reusable receiver coils used internally.

N.M de Souza et al . , .American Journal of Roentgenology, 163 (1994), 607-612 describes, in particular, in vivo imaging of the uterine cervix using a ring-design solenoid receiver coil which is placed intravaginally to envelope the cervix.

Such intravaginal coils have been successfully used at B₀ magnetic field strengths of up to 1.5 T (64 MHz).

In order to transmit the fragile received MR signal, with minimal loss, from the patient to the scanner electronics, it is usual to employ a co-axial cable with 50 ohm characteristic impedance where the driving impedance (in this case from the coil) is also 50 ohms. The series impedance of a 37mm diameter copper loop, typical of an intravaginal ring-design solenoid receiver coil, is only a few milli-ohms. However, by resonating this coil inductance with a suitable capacitance, a much higher impedance (typically greater than 1 kohm) may be generated which can then be transformed down capacitively to the required 50 ohms. Figure 1 shows schematically an MRI detector having a single turn solenoid coil 1 which was used successfully at Bo field strengths of 0.5 T (21 MHz) and 1.5 T (64 MHz) . Tuning capacitor C_t' at the output port of the coil determines the resonance frequency of the coil, giving a correspondingly high dynamic impedance across the coil. The capacitors C_mi and C_m2 are then used to provide the required 50 ohm (resistive) impedance to match the co-axial transmission cable 2.

The coil 1 operates in receive-only mode and must therefore present a high series impedance during the excitation phase of the MR sequence, so as to avoid distortion of the Bi field and heating of local tissue. The blocking impedance is generated by resonating the capacitor C_t' with a short length 3 of coaxial cable, a condition achieved when the PIN diode 4 is forward biased (so that the PIN diode behaves as a low value resistance) . During the signal detection phase, the PIN diode switch 4 is usually reversed biased (so that the PIN diode behaves as a low capacitance parallel plate capacitor) , which removes the blocking impedance.

Although, the coil 1 of Figure 1 was used successfully at 1.5 T (64 MHz), at increased Bo field strengths, operational problems can arise. In particular, the higher resonance frequency associated with the increased field strength requires a reduction in the value of C_t' · However, eventually C_t' becomes comparable to the parasitic capacitance added by the body tissue within the coil. As a result, an increasing proportion of the conservative electric field is intercepted by the lossy body tissue instead of residing in the lumped capacitor C_t' . This causes tissue heating and a corresponding reduction in SNR. This behaviour corresponds to the coil displaying the characteristics of a dipole antenna, with additional loss due to signal radiation. A further operational problem relates to distortion of the Bi excitation field. The solenoid coil 1 of the MRI detector of Figure 1 typically has a short cylindrical structure. For example, a 37 mm diameter, 15 mm deep, copper foil cylinder has been found suitable for cervical imaging. The single turn, short cylindrical structure helps to reduce inductance and hence to reduce conservative electric fields. However, a receive-only coil must also cause minimal disturbance to the uniformity of the Bi excitation field. A typical priority, therefore, is to minimise circulating currents around the receive coil loop during Bi excitation. Nevertheless, as the short cylindrical structure presents a significant cross- sectional area perpendicular to the Bi flux, additional field distortion results, even when the coil loop is open-circuited.

More particularly, if a short cylinder solenoid coil is employed as an external surface coil, the field distortion is usually remote from the volume of interest (VOI) and can be ignored. When employed as an, enveloping receive coil, however, Bi flux may be partially shielded within the VOI, and unacceptable field distortions may result. For example, flip angle accuracy is increasingly important for quantitative studies and advanced imaging pulse sequences, but this accuracy can be degraded by Bi field distortions.

It would be desirable to provide an MRI detector for detecting RF signals in which one or some of the above operational problems are reduced, avoided or overcome.

Accordingly, a first aspect of the present invention provides an MRI detector for detecting RF signals having a solenoid coil which completes a single solenoid turn, the coil being split into a plurality of coaxial sub-coils, each of which has an output port, completes the single solenoid turn and has at least one tuning capacitor positioned therein for tuning the resonance frequency of the sub-coil, the sub-coils being spaced along the axis of the coil and being connected in parallel with each other at their output ports.

By splitting the coil into a plurality of sub-coils, the cross-sectional area of the coil can be decreased, which reduces eddy currents in the coil. In this way, distortion of the Bi excitation field can be significantly reduced.

The RI detector may have any one or, to the extent that they are compatible, any combination of the following optional features .

Typically, in order to reduce eddy currents, the sub-coils are only connected with each other at their output ports.

Typically, e.g. for intravaginal applications, the coil may have two, three, four, five or six sub-coils.

Preferably, each sub-coil has a plurality of tuning capacitors positioned therein which divide the sub-coil into a plurality of sub-coil sections connected in series by the tuning

capacitors. By providing each sub-coil with a plurality of tuning capacitors, for a given coil resonance frequency it is possible to increase the value of the individual tuning capacitors. This allows higher resonance frequencies to be achieved by improving the tuning stability of the coil with body loading, and also by reducing dissipated tissue and radiation losses. Typically, the tuning capacitors are circumferentially spaced around each sub-coil to provide sub- coil sections of about equal length.

The use of a plurality of tuning capacitors is not limited to a solenoid coil which is split into a plurality of sub-coils, but has broader applicability to e.g. non-split coils. Thus, a second aspect of the present invention provides an MRI detector for detecting RF signals having:

a solenoid coil which completes one or more solenoid turns (and preferably a single solenoid turn) , and

a plurality of tuning capacitors positioned in the coil for tuning the resonance frequency of the coil, the tuning capacitors dividing the coil into a plurality of coil sections connected in series by the tuning capacitors. Typically, the tuning capacitors are circumferentially spaced around the coil to provide coil sections of about equal length.

The MRI detector of the first or second aspect may have any one or, to the extent that they are compatible, any

combination of the following optional features.

The coil, or (when the coil is split into a plurality of coaxial sub-coils) each sub-coil, may have at least two, preferably at least four, and more preferably at least six or eight of the tuning capacitors.

Preferably, the MRI detector further has:

a PIN diode positioned in the coil and arranged in series with the tuning capacitors, and

circuitry for controlling the bias of the PIN diode, whereby when the PIN diode is forward biased the coil is able to detect RF signals emitted by spins from a subject in response to an excitation Bi magnetic field, and when the PIN diode is zero or reverse biased the coil is substantially prevented from distorting the excitation Bi magnetic field. Conveniently, the PIN diode can be at an output port of the coil, or (when the coil is split into a plurality of coaxial sub-coils) the PIN diode can be at the output ports of the sub-coils .

If a detector of the type shown in Figure 1, e.g. with the PIN diode positioned outside the coil, were to be modified by positioning a plurality of tuning capacitors in the coil such that the tuning capacitors divided the sub-coil into a plurality of sub-coil sections connected in series by the tuning capacitors, the values of the individual tuning

capacitors could be increased. However, with higher capacitor values, it becomes increasingly difficult to achieve a

sufficient blocking impedance during the excitation phase, as the resonance frequency of the coil is increased. But by positioning the PIN diode in the coil, it is possible to avoid this difficulty. In particular, when zero or reverse biased, the PIN diode behaves as a low value capacitor which does not pass much RF signal due to its high impedance. When forward biased, the PIN diode has, in contrast, a low resistance to RF signals .

The use of a PIN diode in this manner is not limited however to a solenoid coil which is split into a plurality of sub- coils or which has a plurality of tuning capacitors. Thus a third aspect of the present invention provides an MRI detector for detecting RF signals having:

a solenoid coil which completes one or more solenoid turns (and preferably a single solenoid turn) ,

one or more tuning capacitors positioned in the coil for tuning the resonance frequency of the coil,

a PIN diode positioned in the coil (e.g. at an output port of the coil) and arranged in series with the tuning

capacitors, and

circuitry for controlling the bias of the PIN diode, whereby when the PIN diode is forward biased the coil is able to detect RF signals emitted by spins from a subject in response to an excitation Bi magnetic field, and when the PIN diode is zero or reverse biased the coil is substantially prevented from distorting the excitation Bi magnetic field.

The MRI detector of the first or second aspect which has a PIN diode positioned in the coil, or the MRI detector of the third aspect, may have any one or, to the extent that they are compatible, any combination of the following optional

features .

The circuitry for controlling the bias of the PIN diode may comprise an inductor connected in parallel with the PIN diode, the inductor conveniently providing a current path for biasing the PIN diode. However, typically, the zero or reverse biased PIN diode has a residual capacitance. The inductor can therefore also resonate with the residual capacitance of the PIN diode when the PIN diode is zero or reverse biased to produce a blocking impedance which substantially prevents the coil from distorting the excitation Bi magnetic field.

The MRI detector of the first, second or third aspect may have any one or, to the extent that they are compatible, any combination of the following optional features.

The MRI detector may further have:

a transmission cable extending from an output port of the coil, or (when the coil is split into a plurality of coaxial sub-coils) from the output ports of the sub-coils, for transmitting RF signals detected by the coil, and

a further capacitor positioned in the coil at the output port or ports, wherein the further capacitor matches the coil output impedance to the impedance of the transmission cable. The value of the further capacitor does not, typically, have a significant effect on the resonance frequency of the coil. However, the further capacitor does help to avoid the

necessity for additional external capacitors, such as C_mi and C_m2 in the detector of Figure 1, for impedance matching between the coil and the transmission cable.

Typically, the transmission cable is a co-axial cable.

The transmission cable and the further capacitor both

preferably provide an impedance of about 50 ohms at the coil resonance frequency. The solenoid coil may have an internal diameter of less than 40 mm. The solenoid coil may have a depth (i.e. length in the axial direction of the coil) of less than 20 mm.

Preferably, the solenoid coil is tuned to give a resonance frequency for the coil of at least 100 MHz, and more

preferably of at least 120 MHz.

Preferably, the solenoid coil is adapted for internal in vivo use. For example the coil can be contained in a suitable housing which, typically, has to: protect the coil

electronics, have good radiofrequency properties, and be nontoxic, sterilizable and biocompatible. The housing can be formed, for example, of an acetyl homopolymer. The coil may be embedded in epoxy resin for water resistance and to seal the electronics. Preferably the solenoid coil is adapted for vaginal insertion. The coil may be located, for example, at the end of an approximately 30 cm long elongate member which can provide a guide for the transmission cable and can

facilitate positioning of the coil. The solenoid coil can be adapted to be positioned around the cervix, i.e. to envelope the cervix. For such positioning or enveloping, preferably the coil has an internal diameter of about 37 mm and a depth of about 12 mm.

Preferably, the solenoid coil is a receive-only coil.

A further aspect of the invention provides an apparatus for generating a magnetic resonance image of a subject, the apparatus comprising:

a magnet for generating a B₀ magnetic field within the subject,

a transmit coil for generating an excitation Βχ magnetic field within the subject, the Bi magnetic field being

perpendicular to the Bo magnetic field, and

the MRI detector of any one of the first, second and third aspects for detecting RF signals emitted by spins within the subject .

Preferably the magnet produces a B₀ magnetic field having a field strength of at least 3 T.

A further aspect of the invention provides the use of the MRI detector of any one of the first, second and third aspects for internal in vivo imaging of a human or animal body. For example, the use can be in vivo imaging of a cervix.

A further aspect of the invention provides a method of imaging a subject comprising:

inserting the MRI detector of any one of the first, second and third aspects into a human or animal body,

generating a Bo magnetic field within the body,

generating an excitation Bi magnetic field within the body, the Bi magnetic field being perpendicular to the B₀ magnetic field,

detecting RF signals emitted by spins within the body at the detector, and

producing an image of the body from the detected RF signals .

Preferably the Bo magnetic field has a field strength of at least 3 T.

Embodiments of the invention will now be described by way of example with reference to the accompanying drawings in which:

Figure 1 shows schematically an MRI detector having a single turn solenoid coil;

Figure 2 shows results for a hybrid method of moments (HMOM) electromagnetic (EM) simulation of the vertical (y-axis) component of the Bi receive field of a single turn, copper, solenoid coil of the type shown Figure 1; Figure 3 shows results for an HMOM EM simulation of the horizontal (x-axis) component of a Βχ excitation field produced by an excitation loop;

Figure 4 shows results for an HMOM EM simulation of the horizontal (x-axis) component of the Bi excitation field produced by the excitation loop of Figure 3 when the solenoid coil of Figure 2 is positioned therein;

Figure 5 shows results for an HMOM EM simulation of the horizontal (x-axis) component of the Bi excitation field produced by the excitation loop of Figure 3 when a detector coil according to the present invention is positioned therein;

Figure 6 shows results for an HMOM EM simulation of the horizontal (x-axis) component of the Bi excitation field produced by the excitation loop of Figure 3 when a further detector coil according to the present invention is positioned therein;

Figure 7 shows schematically test objects used to load a body coil and cylindrical copper structure;

Figure 8 shows images of the Bi excitation field within the dashed outline in Figure 7 when (a) no cylindrical copper structure was in place, (b) a cylindrical copper structure with a single turn was in place, (c) a cylindrical copper structure with two parallel turns was in place, and (d) a cylindrical copper structure with four parallel turns was in place; and

Figure 9 shows schematically an MRI detector having a single turn solenoid coil with distributed tuning capacitors.

To investigate the problem of distortion of the Bi excitation field, hybrid method of moments (HMOM) electromagnetic (EM) simulations were performed on solenoid coil arrangements. The minor influence of tuning capacitors was not considered in the simulations .

Figure 2 shows results for an H OM EM simulation of the vertical (y-axis) component of the Bi receive field of a single turn, copper, solenoid detector coil 11 of the type shown Figure 1, the gap at the front of the coil corresponding to the position of the coil output port at C_t' . Figure 3 shows results for an HMOM EM simulation of the horizontal (x-axis) component of a Bi excitation field produced by an excitation loop 12. For modelling convenience, a relatively small excitation loop was modelled, which resulted in a less uniform excitation field than would normally be the case. The

excitation loop produced a linear horizontal magnetic field component along the x-axis. Figure 4 then shows results for an HMOM EM simulation of the horizontal (x-axis) component of the Bi excitation field produced by the excitation loop of Figure 3 when the detector coil of Figure 2 is positioned therein. The simulation shows that the eddy currents

generated in the detector coil during Βχ excitation

significantly distort the Bi excitation field, particularly at positions close to the detector coil. Further, there is a reduction in Bi excitation field strength throughout the volume encircled by the detector coil.

The Bi excitation field is typically circularly polarised in the x-y plane, and the orientation of the detector coil 11 ensures that the cross-sectional area presented by the coil perpendicular to the Βχ excitation field varies throughout each cycle of the Bi excitation field, being at a maximum with Bi in the x direction and at a minimum with Βχ in the y direction. Faraday's Law dictates that any reduction in the cross- sectional area of the copper presented perpendicular to Bi commensurately will reduce eddy currents. To achieve such a reduction in cross-sectional area in a detector coil according to the present invention, the coil was split into two coaxial sub-coils. Figure 5 shows results for an HMOM EM simulation of the horizontal (x-axis) component of the Bi excitation field produced by the excitation loop of Figure 3 when such a detector coil 21 is positioned therein. The coaxial sub-coils are spaced from each other along the y- axis, and both complete the single solenoid turn of the detector coil 11 of Figures 2 and 4. However, as the depth of the detector coil 21 in the y direction is not increased, the overall cross-sectional area in the x direction is reduced. Relative to the detector coil 11, this leads to a decrease in the distortion of the Bi excitation field, and also to an increase in the field strength throughout the volume encircled by the detector coil.

Further improvements in performance can be obtained by

splitting the coil into more sub-coils. Figure 6 shows results for an HMOM EM simulation of the horizontal (x-axis) component of the Bi excitation field produced by the excitation loop of Figure 3 when such a detector coil 31 is positioned therein. The detector coil 31 now has four coaxial sub-coils spaced from each other along the y-axis, and again the depth of the detector coil 31 in the y direction is not increased.

To experimentally verify the simulations, three short

cylindrical copper structures were built on 37mm diameter x 12mm deep acetyl formers, the structures consisting of, respectively, one, two and four parallel turns. The

structures were "open circuits" and therefore presented a high impedance at the resonant frequency, as is desirable for receive-only coils during RF excitation. The coils were tested at 3T in an MRI scanner by measuring the Bi excitation field of the body coil, with and without the cylindrical structures in the excitation RF field. As shown schematically in Figure 7, a large test object 42 was used to provide a suitable load to the body coil and a small cylindrical test object 43 placed on the large test object was used to load the cylindrical structures 41. Bi measurements were made using a sequence with two interleaved repetition times (30 and 150 ms) .

Figure 8 shows images of the Bi excitation field within the dashed outline of Figure 7 when (a) no cylindrical copper structure was in place, (b) the "unsplit" cylindrical copper structure with a single turn was in place, (c) the "split" cylindrical copper structure with two parallel turns was in place, and (d) the "split" cylindrical copper structure with four parallel turns was in place. On introduction of the single-turn, 12mm deep, open-circuit, receive-only coil

(Figure 8(b)), there was a significant reduction in field throughout the encircled volume, the central Βχ field dropping to 89%. Use of two turns (Figure 8(c)), each of 4.5mm depth, led to a marked improvement in field distortion, the central Βχ field only dropping to 99%, although some fall-off is still apparent close to the cylindrical structure. Finally, use of four turns (Figure 8(d)) resulted in barely discernible field disturbance .

Turning to the operational problem in the MRI detector of Figure 1 of tissue heating and SNR reduction at higher B₀ field strengths, Figure 9 shows schematically an MRI detector according to the present invention having a single turn solenoid coil 51 in which, instead of a single tuning

capacitor C_t bridging the ends of the coil at the output port thereof, N tuning capacitors C_t are circumferentially

distributed around the coil with a further impedance-matching capacitor C_m bridging the ends of the coil at the output port. The tuning capacitors thus divide the coil into coil sections 52. Each of these coil sections has an inductance which is approximately 1/(N+1) the inductance of the equivalent

undivided coil (i.e. coil 1 of Figure 1) and is resonated in series with the tuning capacitors. To provide a resonance frequency for the divided coil 51 which is approximately the same as that of the equivalent undivided coil, each tuning capacitor has a value that is N times the capacitance of the single tuning capacitor of the equivalent undivided coil.

Thus, by dividing the coil 51 into coil sections, the values of the tuning capacitors can be significantly increased, such that the tuning stability of the coil 51 with body loading is substantially improved. For example, the coil 51 can be tuned to provide a resonance at 3 T (128 MHz), well above the 1.5 T (64 MHz) achieved with the coil 1 of Figure 1. Dissipated tissue and radiation losses can also be reduced.

Although the solenoid coil 51 is shown as an unsplit coil, the principle of dividing the coil into coil sections can be applied to the split coils discussed above in relation to Figures 5 to 8. In this case, each sub-coil can be divided into sub-coil sections by respective tuning capacitors

circumferentially distributed around the sub-coil.

In the coil 1 of Figure 1, a high series impedance during the excitation phase of the MR sequence is achieved by forward biasing the PIN diode 4 so that the capacitor C_t' at the output port of the coil resonates with the short length 3 of coaxial cable to generate a blocking impedance. The performance of this circuit is dependent, however, on the physically

realisable magnitude of the blocking impedance. In the MRI detector of Figure 9, the capacitor C_m is positioned at the output port of the coil 51. As the resonance frequency of the coil 51 is determined by the tuning capacitors C_t, the value of C_m can conveniently be set to match the, typically 50 ohm, impedance of co-axial transmission cable 53 which extends from the output port. However, if the approach to producing a blocking impedance used in respect of the detector of Figure 1 is adopted for the detector of Figure 9, it becomes

increasingly difficult when the value of C_m is determined by impedance-matching requirements to achieve a sufficient blocking impedance during the excitation phase as the

resonance frequency of the coil 51 is increased.

To overcome this difficulty, in the MRI detector of Figure 9, a PIN diode 54 is positioned in the coil 51, in series with the tuning capacitors C_t and impedance-matching capacitor C_m, rather than being positioned outside the coil as in the detector of Figure 1. Further, an inductor L is connected in parallel across the PIN diode 54 and capacitor C_m to provide a current path for the diode d.c. drive.

During the excitation phase of the MR sequence, the PIN diode 54 is zero or reversed biased. One option is to choose the inductor L so that it is close to self-resonance at the resonance frequency of the coil 51 (i.e. about 128 MHz at 3 T) . However, in practice, the zero or reversed biased PIN diode has a very low residual capacitance, typically of about 2.2 pF, which gives an impedance of about 566 ohms at 128 MHz. Therefore, preferably the inductor L is chosen to have a similar impedance (e.g. about 566 ohms at 128 MHz) to the zero or reversed biased PIN diode so that the inductor resonates with this residual capacitance. In this way, impedances of about 15 kohms in series with the coil 51 between positions A and Ai can be achieved, which provides effective detuning of the coil.

A preamplifier input at the other end of transmission cable 53 is typically protected by shunting its input with a protecting PIN diode during the excitation phase. This causes the impedance "seen" at the coil end of the transmission line to depend on the line length. By adjusting the line length to be a half-wavelength (approximately 0.775 metres at 128 MHz), a low resistance is presented across C_m, which avoids dependency of the detuning circuit on the value of Cm.

During the receive phase, the PIN diode 54 is forward biased to provide a low value resistance- The presence of the inductor L (providing an impedance of e.g. about 566 ohms) connected across the 50 ohm output port has little influence on circuit performance.

Although there are losses at the PIN diode 54 during signal detection, these losses are relatively insignificant compared to the unavoidable body losses due to non-conservative electric fields at high frequencies such as 128 MHz.

As mentioned above, the inductor L can be chosen so as to resonate with the capacitance of the PIN diode 54 when not forward-biassed. This maximises the impedance of the blocking circuit when in the excitation phase. In this resonant condition, however, the r.f. current circulating within L is increased, and significant magnetic field can exist around L. This field can distort the local Bi excitation field, leading to image artefacts. This problem can be reduced or overcome by positioning L away from the imaging field of view within, e.g. along the coil "handle" which also carries the co-axial transmission cable 53. Additionally or alternatively, a conductive shield may be employed around L.

The Bi excitation field produced by the transmit coil has both electrostatic and magnetic components. Although not directly involved in the MR imaging process, the electrostatic

component (usually known as the E-field) generally should not be ignored, as there may be associated safety implications.

A conductor running parallel to the E-field can act as short- circuit which may carry a potentially large "common mode" current. This can be the case when a cable connecting a receive-only coil passes between a patient and scanner input electronics outside the magnet bore. At frequencies at or above 100 MHz, such cables may have a physical length close to a half wavelength and a resonance can occur, amplifying the induced r.f. current and increasing the hazard. An insertable MR probe of the type shown schematically in Figure 9 (adapted to be positioned around the cervix), may be more subject to the effects of these induced r.f. currents (compared to a standard surface coil) due to stronger capacitive coupling with the body.

However, the problem can be addressed by introducing

impedances in series with the connecting cable 53 which attenuate the "common mode" current whilst not affecting the fragile received signal (i.e. the "differential mode"). As well as reducing the induced current in general, a number of these circuits, often referred to as Baluns, may be

strategically positioned along the cable so as to help prevent resonances. In the case of the insertable cervix coil of Figure 9, a minimum of two baluns are generally employed, one close to the coil, within the handle, so as to provide a common mode impedance between the cable and the patient, and a second balun along the cable approximately one quarter wavelength from the first. See D.M. Peterson et al. Common Mode Signal Rejection Methods for MRI: Reduction of Cable Shield Currents for High Static Magnetic Field Systems,

Concepts in Magnetic Resonance Part B (Magnetic Resonance Engineering), 19B(1) (2003), 1-8.

While the invention has been described in conjunction with the exemplary embodiments described above, many equivalent modifications and variations will be apparent to those skilled in the art when given this disclosure. Accordingly, the exemplary embodiments of the invention set forth above are considered to be illustrative and not limiting. Various changes to the described embodiments may be made without departing from the spirit and scope of the invention.

All references referred to above are hereby incorporated reference .

Claims

1. An MRI detector for detecting RF signals having:

a solenoid coil (51) which completes one or more solenoid turns ,

one or more tuning capacitors (C_t) positioned in the coil for tuning the resonance frequency of the coil,

a PIN diode (54) positioned in the coil and arranged in series with the tuning capacitors, and

circuitry for controlling the bias of the PIN diode, whereby when the PIN diode is forward biased the coil is able to detect RF signals emitted by spins from a subject in response to an excitation Bi magnetic field, and when the PIN diode is zero or reverse biased the coil is substantially prevented from distorting the excitation Βχ magnetic field.

2. An MRI detector according to claim 1, wherein the coil completes a single solenoid turn.

3. An MRI detector according to claim 2, wherein the coil is split into a plurality of coaxial sub-coils, each of which has an output port, completes the single solenoid turn and has at least one tuning capacitor positioned therein for tuning the resonance frequency of the sub-coil, the sub-coils being spaced along the axis of the coil and being connected in parallel with each other at their output ports.

4. An MRI detector according to any one of the previous claims, wherein the circuitry comprises an inductor (L) connected in parallel with the PIN diode, the inductor

providing a current path for biasing the PIN diode.

5. An MRI detector according to claim 4, wherein the inductor resonates with a residual capacitance of the PIN diode when the PIN diode is zero or reverse biased to produce a blocking impedance which substantially prevents the coil from

distorting the excitation Bi magnetic field.

6. An MRI detector according to any one of the previous claims, wherein the solenoid coil is a receive-only coil.

7. An MRI detector for detecting RF signals having a receive- only solenoid coil (21, 31) which completes a single solenoid turn, the coil being split into a plurality of coaxial sub- coils, each of which has an output port, completes the single solenoid turn and has at least one tuning capacitor positioned therein for tuning the resonance frequency of the sub-coil, the sub-coils being spaced along the axis of the coil and being connected in parallel with each other at their output ports .

8. An MRI detector according to claim 3 or 7, or according to claim 4, 5 or 6 as dependent on claim 3, wherein each sub-coil has a plurality of tuning capacitors positioned therein which divide the sub-coil into a plurality of sub-coil sections connected in series by the tuning capacitors.

9. An MRI detector according to claim 1 or 2, or according to claim 4, 5 or 6 as dependent on claim 1 or 2, wherein the coil has a plurality of tuning capacitors positioned therein which divide the coil into a plurality of coil sections connected in series by the tuning capacitors.

10. An MRI detector according to any one of the previous claims, further having:

a transmission cable (53) extending from an output port of the coil, or from the output ports of the sub-coils, for transmitting RF signals detected by the coil, and

a further capacitor (C_m) positioned in the coil at the output port or ports, wherein the further capacitor matches the coil output impedance to the impedance of the transmission cable .

11. An MRI detector according to any one of the previous claims, wherein the solenoid coil has an internal diameter of less than 40 mm.

12. An MRI detector according to any one of the previous claims, wherein the solenoid coil is tuned to give a resonance frequency for the solenoid coil of at least 100 MHz.

13. An MRI detector according to any one of the previous claims, wherein the solenoid coil is adapted for internal in vivo use.

14. An MRI detector according to claim 13, wherein the solenoid coil is adapted for vaginal insertion.

15. An MRI detector according to claim 13 or 14, wherein the solenoid coil is adapted to be positioned around the cervix.

16. An apparatus for generating a magnetic resonance image of a subject, the apparatus comprising:

a magnet for generating a Bo magnetic field within the subj ect ,

perpendicular to the Bo magnetic field, and

the MRI detector of any one of the previous claims for detecting RF signals emitted by spins within the subject.

17. Use of the MRI detector of any one of claims 1 to 15 for internal in vivo imaging of a human or animal body.

18. Use according to claim 17 for in vivo imaging of a cervix.