WO2010018479A1

WO2010018479A1 - Magnetic resonance rf coil

Info

Publication number: WO2010018479A1
Application number: PCT/IB2009/053202
Authority: WO
Inventors: Cornelis W. Jacobs
Original assignee: Koninklijke Philips Electronics N.V.
Priority date: 2008-08-13
Filing date: 2009-07-23
Publication date: 2010-02-18

Abstract

The invention relates to a nuclear magnetic resonance coil comprising electrically conductive coil loop segments (100; 102), wherein the coil loop segments are alternately arranged on opposed sides of a dielectric material layer (202), wherein consecutively arranged coil loop segments partially overlapping each other.

Description

Magnetic resonance RF coil

TECHNICAL FIELD

The invention relates to a nuclear magnetic resonance coil, a method of design and manufacturing a nuclear magnetic resonance coil and a computer program product.

BACKGROUND AND RELATED ART

Magnetic resonance imaging (MRI) is a state of the art imaging technology which allows cross-sectional viewing of objects like the human body with unprecedented tissue contrast. MRI is based on the principles of nuclear magnetic resonance (NMR), a spectroscopic technique used by scientists to obtain microscopic chemical and physical information about molecules. The basis of both NMR and MRI is the fact, that atomic nuclei with non-zero spin have a magnetic moment. In medical imaging, usually nuclei of hydrogen atoms are studied since they are present in the body in high concentrations like for example water. The nuclear spin of elementary particles can resonate at the resonance frequency, if a strong DC magnetic field is applied. This magnet resonance (MR) frequency is determined by the level of the magnetic flux. In an MRI scanner, the magnetic field matches a selected resonance frequency only at one position in space. Only at this position the presence of these particles can be detected. By varying this position step by step, an image can be measured. In practice, more sophisticated algorithms are used to achieve the image in a reasonable time from e.g. 'slices' of the investigated volume. The slice section is derived with help of gradient coils. Typical resonance frequencies are in the range from 40 MHz to 300 MHz, corresponding to magnetic flux levels in the range of 1 T to 7 T.

The needed strong DC magnetic field (Bo field) is typically generated by superconducting magnets. In order to vary this field, such that it matches a given radio frequency only at one position, a field gradient is generated using gradient coils. The field gradient can vary over time to achieve this gain. The frequency range in the gradient coils is low, and reaches up to a maximum of 10 kHz.

To excite nuclear resonance, an RF coil generates a high frequency magnetic field (Bi field) at the nuclear resonance. This high frequency magnetic field is orientated perpendicular to the strong DC magnetic field. To measure nuclear resonances, receiver coils are placed close to the region of interest, e.g. on a patient. Often a number of sensor coils are connected to a complete module, e.g. such a module may consist of 12 (= 3 x 4), 16 (= 4 x 4) or 24 (= 4 x 6) individual sensor coils, which is called a coil array. The module also includes additional electronics to process the measured signals or switching between transmit (Tx) and receive (Rx) mode.

The receiver RF coil can be either the transmit coil itself or an independent receive only RF coil. In case of a receive only RF coil (array) it has to be detuned during the transmit phase.

The signal to noise ratio (SNR) and spectra resolution of acquired MR data are increasing with increasing magnetic field strength. Since the resonance frequency of nuclei of investigated atoms varies as a function of the applied magnetic field, an RF coil needs to be tuned to a certain resonance frequency to be power efficient in excitation of these nuclei and/or in order to be sensitive to these RF frequencies. Typical hydrogen nuclear frequencies are approximately 64 MHz for a field strength of 1.5 T, 128 MHz at 3T and 300 MHz at 7 T. A proper tuning of the resonance frequency of coils can for example be performed by including a capacitor into the coil loop. However, with increasing RF frequency, the required capacitance for tuning such a resonance circuit is significantly decreased, which in turn leads to the disadvantage, that such a capacitor with low capacitance is subject to extremely high voltages in case of transmit coil (arrays). In order to overcome this problem, typically the voltage over such a tuning capacitor is divided over a set of discreet capacitors in series distributed over the coils geometry. This also helps to decrease the electrical coupling with the patient since the induced high voltages are thus divided in some lower voltages in series. In practice this is done by splitting the capacitor into more capacitors in series divided over the coils geometry. Further, a coil must be optimized regarding coil losses which may occur due to the finite resistance, i.e. due to the skin effect of the coil conductors and the ohmic losses in the HQ multilayer ceramic capacitors used to tune the RF coil. Conductors are preferred to be relative small to avoid too much capacitance (electrical coupling) to the examined object like a patient as well as for reason of eddy currents induced by fast switching gradient fields. Usually, the coil quality factor Q which is inverse proportional with the RF losses is used while designing and evaluating the coil. The goal is a high unloaded coil quality factor Q_u because this indicates that a coil has minimal coil losses. Furthermore, the coupling of the patient with the E-field has to be minimized. Typically an optimization with respect to these coil properties is also performed by using discreet capacitors distributed in series over the coils geometry.

Preferably, the geometry, i.e. the shape and the dimensions of the coil are determined mainly by the required penetration depth and the region of interest inside the patient's body. The higher her frequencies are the smaller the required capacitor value to tune the coil becomes. As a consequence, typically specific selected capacitor values are required and / or capacitors need to be fine tune to properly tune the coil.

However, the usage of discreet capacitors has the disadvantage, that an expensive soldering of multilayer capacitors with discreet values between separate coil segments is required, which is expensive and also due to the presence of a large amount of soldering points error prone. Further, in case coil arrays are used, individual coils of the coil array need to be decoupled which is typically performed by either partially overlapping the coils or interconnecting the coils by further capacitors. However, this often leads to rather chaotic situations in which a large amount of discreet capacitors need to be arranged together within a coil unit which often is space consuming and does not allow to build slim- shaped RF coil units. Further, in case bended RF coils are built, individual discreet capacitors need to be connected to the RF coil circuit by hand, since it is rather difficult to automatically place discreet components to conductors on bended surfaces.

It has to be noted here, that a coil segment is the piece of a coil loop (RF antenna) between the capacitors that are used to tune the coil loop.

US 2003/0048103 Al discloses an increased distributed capacitance for an NMR resonant probe coil which is obtained by a surrounding conducting surface which is congruent to and spaced from the major portion of the RF current density distribution on a surface on the inductive component of the resonator. There is a need for an improved nuclear magnetic resonance coil, as well as an improved method of designing and manufacturing a nuclear magnetic resonance coil.

SUMMARY OF THE INVENTION

The present invention provides a nuclear magnetic resonance coil comprising electrically conductive coil loop segments, wherein the coil loop segments are alternately arranged on opposed sides of a dielectric material layer, wherein consecutively arranged coil loop segments are partially overlapping each other.

The coil according to the invention is applicable for magnetic resonance coils used for transmit and/or receive purpose. It can be used in MRI as a transmit and/or receive coil, or in NMR. By means of the coil geometry in which a capacitance between individual coil loop segments is formed by said coil loop segments and the dielectric material layer allows avoiding the need for individual discreet capacitors. Preferably, the dielectric material layer comprises a high frequency laminate material. By balancing the overlapping and/or width of the coil loop segments, in combination with the thickness and 8_R of the dielectric material, almost any embedded capacitance can be realized. Since the physical length of the capacitor can be individually adapted, the capacitance can be equally distributed over the actual coil geometry.

Further, by such an arrangement of coil loop segments, it is possible to realize serially connected capacitors in an alternating manner.

In accordance with an embodiment of the invention, the coil loop segments have the same segment length, wherein the consecutively arranged coil loop segments are overlapping each other in equal shares. In other words, the length of the overlapping coil loop segments that make a capacitance is preferably to be chosen about the same length of the coil loop segment that is connecting it to the next capacitance in an RF loop. This way, the dimension (or the length) over which a voltage due to RF currents is built up is equal to the length of the capacitance inverting this voltage. This spreads the induced voltage more equally over the coil element lowering the E-field coupling with the patient. Further, by building large capacitors using overlapping areas and high-frequency laminate materials as dielectric materials allow to avoid to solder many small and less efficient individual discreet capacitors. As a consequence, a reduced number of soldering points provides a reduction in resistive coil losses.

Further, for example when the coil is used to transmit a high-power transmit field, the induced voltages are better spread over the whole coil geometry. The voltage is not concentrated at the relatively small dimensions of a discreet capacitor. Thus, emerging resistive heat is not concentrated in hotspots like for example small ceramic multilayer capacitors and soldering points but the heat is equally distributed over the large overlapping area. The avoidance of individual hotspots reduces for example the burning risk of a patient imaged by means of such a coil or in NMR it reduces the risk of locally overheating certain substance areas investigated by such a coil. Also, avoiding of hotspots allows preventing an affecting of the dielectric properties, for example the dielectric losses (loss tangent), of the capacitors.

Further, because there is no need to solder discreet capacitors in the coil, less discreet components and production steps will result in more reliability. In accordance with a further embodiment of the invention, the coil comprises a coil support with top and a bottom side, wherein the coil loop segments are carried by the coil support on the top and bottom side, wherein the dielectric material layer is formed by the coil support itself. This allows to manufacture such a coil in a machine made manner, since for example simple patterning techniques followed by patterning and coating as known from printed circuit board production procedures can be directly applied to the coil support without the need to further apply many discreet expensive high Q multilayer ceramic capacitors. Especially when applied in coil arrays this becomes very effective. There is further no need for additional rigid boards carrying the discreet capacitors to allow the RF loop to be bended without breaking the discreet capacitors. For example, as coil support a flexible dielectric material can be used, on which by standard printed circuit board (PCB) manufacturing techniques conducting layers for example made of thin copper foil can be applied. Afterwards, the flexible substrate can be bended into almost any suitable form in order to be used as a magnetic resonance coil. The usage of individual discreet capacitors would require manufacturing the bended coil structure followed by manually attaching the individual capacitors. As a consequence, the RF coil would always need to be kept at its bended structure without the possibility to individually adjust the bending adapted for individual imaging purposes on different patients. However, in case a flexible dielectric material is used, due to the thin and thus highly flexible coil loop segments an individual adaption and forming of the RF coil according to the invention is possible without any risk of breaking capacitors and thus damaging the RF coil.

In accordance with an embodiment of the invention, the coil comprises a primary magnetic field direction, wherein the coil loop segments have a planar structure, wherein the perpendicular of the planar structure is parallel to the magnetic field direction. Typically, this magnetic field direction is indicated by Bi. In case of such a 'horizontal' arrangement of the coil loop segments with respect to the primary magnetic field direction, there is no additional space requirement for discreet multilayer capacitors or space required to mount additional boards which carry these discreet components as known in the art. Thus, such RF coils can be made very thin. Further, the RF coil arranged in horizontal orientation has the advantage of less coupling of the capacitance with the patient's body. This is due to the fact, that the coil loop segments further have a shielding effect on the voltage across the capacitance formed by the coil loop segments used in the RF coil.

In accordance with a further embodiment of the invention, the coil comprises a primary magnetic field direction, wherein the coil loop segments have a planar structure, wherein the perpendicular of the planar structure is orthogonal to the primary magnetic field direction. This is also called the 'vertical orientation' of the coil loop segments. The usage of such a vertical orientation has the advantage, that resistive conductive losses in the coil loop segments are minimal. This is due to the fact, that at higher RF frequencies the skin effect leads to a current conduction only on the surface of the coil loop segments. For example, in case copper is used as coil loop segment material, the skin depth at 64 MHz is only 8.2 μm, which is further decreased for even higher frequencies. Since in the vertical orientation the coil loop segments are in parallel with the orientation of the Bi field, resistive losses are minimal as there is more copper or conductive area available to conduct the current. Further, because the coil loop segment area facing the patient is smaller, the capacitance between the patient and the RF coil is minimized.

Preferably, the thickness of the planar coil loop segments is no more than 17 μm. This provides a good mechanical stability of the coil loop segments and also attenuates low frequency Eddy currents induced by gradient coils which would slow down the gradient field switching frequency due to counteracting gradient fields emerging from the Eddy currents.

It has to be noted here, that 'planar structure' has to be understood as a spatial structure which has the shape of a ribbon, but which itself can be bended.

In accordance with an embodiment of the invention, the coil is a printed circuit board. In other words, preferably the printed circuit board body acts as the dielectric material layer which carries further conducting layers which are made for example of thin copper foils. As printed circuit board based materials many kinds of high-frequency laminate materials can be used. The dielectric properties of such high-frequency laminate materials are comparable or even better to the dielectric properties of high Q multilayer ceramic capacitors and much better than the current base materials used to carry the high Q multilayer ceramic capacitors, wherein these materials include like for example FR-4 (woven glass and epoxy) materials. Further, the PCB body material can be chosen to be formable to follow the optimal shape for patient comfort and to fit the patient anatomy as the capacitances can follow the exact shape of the RF loop and can also have any arbitrary shape. One preferred material is for example Rogers 3003.

Further, by applying state of the art printed circuit board manufacturing techniques, required capacitances can be achieved by appropriate design of the boards in a highly exact manner without the need to compromise for discreet steps or additional discreet fine-tune capacitors in parallel. Also the location where the capacitance is located on the board can be chosen freely without compromising for mechanical interferences known when applying individual discreet capacitors.

In another aspect, the invention relates to a nuclear magnetic resonance coil array comprising a first coil according to the invention and a second coil according to the invention. A reason for combining individual coils is to increase the SNR by replacing a single coil with an array of smaller coils and to use multiple receivers to add the signals together at the image construction stage. However, the coils in such a coil array must be decoupled to prevent signals from one coil interfering with signals from another coil. According to an embodiment of the invention, the first coil is spatially separated from the second coil, wherein the first and the second coil are decoupled from each other by a decoupling capacitance, wherein the decoupling capacitance is formed by a first and a second conductive area and a dielectric decoupling material layer, wherein the first conductive area is electrically connected to a coil segment of the first coil and wherein the second conductive area is electrically connected to a coil segment of the second coil, wherein the first conductive area is spatially overlapping the second conductive area in a further overlapping region, wherein the dielectric decoupling material is located in the further overlapping region. Preferably, the first conductive area is carried by the coil support on the top side and the second conductive area is carried by the coil support on the bottom side, wherein the dielectric decoupling material layer is formed by the coil support. The embedded capacitor between those two loops is taking care for the decoupling of the two loops as the capacitive coupling is counteracting for the inductive coupling of the two loops that are close to each other.

In accordance with another embodiment of the invention, the first and the second coil are overlapping for a mutual electromagnetically decoupling of the first and the second coil, wherein the first and the second coil are located in different layers of a multilayer printed circuit board.

In another aspect, the invention relates to a method of manufacturing a nuclear magnetic resonance coil comprising electrically conductive coil loop segments, wherein the coil loop segments are alternately arranged on opposed sides of a dielectric material layer, wherein consecutively arranged coil loop segments are partially overlapping each other.

In another aspect, the invention relates to a computer program product comprising computer executable instructions to perform any of the method steps of the method of manufacturing a nuclear magnetic resonance coil according to the invention. Thus, by means of the coil according to the invention it is possible to split up capacitances into more capacitors in series to divide high voltages (TX) and lower the patient - coupling effects (RX) that lower the SNR and cause frequency shift of the antenna. The required capacitors can follow the shape of the optimal antenna-loop and be divided into as many as required capacitors without the need to compromise with the mechanical support.

BRIEF DESCRIPTION OF THE DRAWINGS

In the following preferred embodiments of the invention are described in greater detail by way of example only making reference to the drawings in which: Fig. Ia is a schematic of state of the art coil loop segments connected to each other by a capacitor,

Fig. Ib is a schematic of state of the art coil loop segments connected to each other whereby the capacitance ice divided into two (discreet) capacitors,

Fig. 2a is a schematic illustrating coil loop segments coupled to each other by a capacitance according to the invention,

Fig. 2b is a schematic illustrating coil loop segments coupled to each other by two capacitances according to the invention,

Fig. 3 is a schematic illustrating a prior art coil arrangement,

Fig. 4 is a schematic illustrating a coil arrangement according to the invention, Fig. 5 is a prior art coil arrangement in vertical orientation,

Fig. 6 is a schematic illustrating a coil arrangement according to the invention in vertical orientation,

Fig. 7 is a schematic illustrating a prior art coil array,

Fig. 8 is a schematic illustrating a coil array according to the invention, Fig. 9 illustrates the voltage built up in a magnetic resonance coil according to the invention,

Fig. 10 is a schematic illustrating a nuclear magnetic resonance coil comprising a set of capacitances C, inductances L and resistances R,

Fig. 11 is a schematic illustrating a nuclear magnetic resonance coil considering additionally a series resistance typically caused by the patient loading of the coil,

Fig. 12 illustrates a capacitive tap,

Fig. 13 depicts another embodiment in which a high input impedance preamp is used. DETAILED DESCRIPTION

In the following similar elements are depicted by the same reference numerals. Fig. 1 is a schematic illustrating a prior art coil loop section which comprises a first coil loop segment 100 and a second coil loop segment 102. The first and the second coil loop segments are connected to each other by means of a capacitor 108 which comprises the active capacitor material 104 itself as well as soldering points 106 interconnecting the capacitor material 104 to the first coil loop segment 100 and the second coil loop segment 102.

As described above, according to the invention this capacitor 108 is substituted by means of an overlapping area in the layout of the coil segments with a dielectric material layer in between, which is illustrated in fig. 2a. A second coil loop segment 102 is partially overlapping a first coil loop segment 100. Preferably, the dielectric material layer 202 is for example the base material of a printed circuit board, wherein the coil loop segments 100 and 102 are printed on the front and backside, respectively, of said base part 202 of the circuit board.

A single capacitor is formed by the first coil loop segment 100, the dielectric material layer 202 and the second coil loop segment 102. However, by alternately stringing together first and second coil loop segments, a serially connected capacitor arrangement can be obtained. This is illustrated in Fig. 2b, in which alternately first (100) and second (102) coil loop segments are arranged spatially separated on each side of the dielectric material layer 202.

Fig. 3 is a schematic illustrating a prior art arrangement of capacitors arranged in series over a coil geometry. The coil comprises two connection points 300, which may be for example connected with a high-power transmitter and/or a preamplifier (preamp). Individual coil loop segments 100 and 102 are located on a support 302 and by means of capacitors 108 interposed between individual coil loop segments 100 and 102 the coupling of the E-field with the patient is minimized. The induced voltage is divided with the capacitors 108.

Further, the coil loop segments are arranged in such a manner, that the ribbon- like coil loop segments are arranged orthogonal with respect to the orientation of the primary magnetic field direction 304 associated with the coil depicted in fig. 3.

Fig. 4 illustrates the same arrangement of coil loop segments with respect to the primary magnetic field direction, however for a nuclear magnetic resonance coil according to the invention. Here, the first coil loop segments 100 and the second coil loop segments 102 are alternately arranged on opposed sides of the dielectric material layer 202 which on its top and bottom sides carries the coil loop segments. The first and second coil loop segments are partially overlapping each other. In other words, by means of the dielectric material layer 202 and the coil loop segments 100 and 102, the capacitors 108 shown in the prior art as indicated in fig. 3 are substituted.

Preferably, in fig. 4 no additional substrate material carrying the coil loop segments and the dielectric material layer is required, since the dielectric material layer is the substrate itself. In other words, preferably a circuit board made of a high-frequency laminate material is printed on its top and bottom with the coil loop segments 100 and 102. Fig. 5 illustrates a schematic of a prior art RF coil in 'vertical orientation'. In other words, the coil loop segments 100 and 102 are arranged in such a manner, that the planar structure of these ribbon- like coil loop segments is oriented parallel to the primary magnetic field direction 304 of the coil. As a consequence, the capacitors 108 interconnecting the individual coil loop elements 100 and 102 are sticking out from the coil surface in radial direction of the circular coil depicted in fig. 5.

In contrast, as depicted in fig. 6 according to the invention the capacitors formed by means of the coil loop segments 100 and 102, as well as the dielectric material layer 202 are an integral part of the coil structure itself. The planar structure of the coil loop segments 100 and 102 extend parallel to the primary magnetic field direction 304 such that resistive losses will be minimal as there is more copper area available to conduct the current. The RF current follows the path of least self- inductance. Further, the coil loop segment area facing the object to be imaged, for example the patient, is smaller such that the capacitance between the patient and the RF coil is minimized which as a consequence also minimizes the capacitive coupling of the object to be imaged with the coil element. It has to be mentioned here, that preferably the object to be imaged is located in a plane oriented perpendicularly to the magnetic field direction 304 and located below or above the coils depicted in figs. 3-6. However, in general the coil geometry can follow the optimal shape of the object to be imaged. This may require for example that the dielectric material layer and the ribbon like conductors are flexible to be accordingly bended. Fig. 7 is a schematic illustrating a prior art coil array consisting of a first coil

700 and a second coil 702. Both coils comprise discreet capacitor elements 108 which are interposed between coil loop segments 100 and 102. The coils 700 and 702 are partially overlapping to cancel mutual inductance. In case the coils would not be overlapping, due to the circular run of close magnetic field lines one coil would always be penetrated by magnetic field lines generated by the other coil in such a manner, that the magnetic field lines always point only in one given direction for that coil. Thus, continuously a current may be induced due to the field lines which current may be disturbing for said coil. By overlapping the first and the second coils, magnetic field lines induced by the first coil are penetrating through the second coil in the overlapping region in a first direction, whereas outside the overlapping region the magnetic field lines generated by the first coil are penetrating through the second coil in a direction opposite to the mentioned penetration direction. Thus, induced currents are compensated within the second coil.

Instead of using individual discreet capacitors 108, in fig. 8 capacitors are used which are formed by the first and second coil loop segments 100 and 102, respectively, as well as the dielectric material layer 202. As depicted in fig. 8, two coils 804 and 806 comprising such kinds of capacitors are arranged adjacent to each other in a non-overlapping, i.e. spatially separated manner in a coil array, wherein the coils 804 and 806 are decoupled using a decoupling capacitor 800 and 802. The capacitances 800 and 802 are each formed by a first conductive area 808 and a second conductive area 810 and the dielectric decoupling material layer 202, wherein the first conductive area 808 is electrically connected to the coil segment 102 of the coil 804 and wherein the second conductive area 810 is electrically connected to the coil segment 202 of the second coil 806. The first conductive area 808 is partially overlapping the second conductive area 810 in an overlapping region, wherein the dielectric decoupling material 202 is located in this overlapping region.

Again, preferably the coils 804 and 806 are both formed on a common substrate as a printed circuit board, wherein the substrate of this board is a high-frequency laminate material. Further, preferably the first and the second conductive areas 808 and 810 are carried by the dielectric material 202 on its opposite sides. It has to be noted here, that compared to overlapping coil array arrangements, the coil arrangement depicted in fig. 8, has the advantage, that for example for imaging using the SENSE-technique the individual coils are more easily distinguishable with respect to their sensitivity profiles. As a consequence, by means of the coil array depicted in fig. 8 SENSE imaging is made more efficient. In a further embodiment of the invention not depicted here, the coils 804 and

806 may also be arranged mutually overlapping as depicted in fig. 7, however with the capacitors arrangement 100 and 102, wherein the coils 804 and 806 are located in different layers stacked above each other in different layers of a multilayer printed circuit board. Fig. 9 illustrates the voltage built up in a coil according to the invention. The coil with its coil segments 100 and 102 can be subdivided into different parts having more inductive (L) or more capacitive (C) character. As mentioned above, preferably the length of the coil segment overlapping area that makes the capacitance is best to be chosen about the same as the length of the segments that are connecting the overlapping areas. This way, the dimension (length) over which the voltage is built up is equal to the length of the capacitance inverting this voltage. This spreads the induced voltage more equally over the whole coil thus lowering the E-field coupling with the patient.

As the total coil loop dimension is determined by the region of interest and the patient shape or the shape of the object to be imaged, the inductance of such kind of loop is always fixed. Thus, in order to make this loop resonating, the capacitance needs to be accordingly adjusted. The higher the frequency, the smaller is the capacitance needed to make the coil resonating. For example, in case a coil is used which has a diameter of around 127mm, and which is subdivided into individual coil loop segments of a length of each 25mm, the inductance can roughly be calculated to 14OnH. In order to make this coil resonating at a system frequency of 63.864 MHz (corresponding to a 1.5T magnetic field), a total capacitance of the coil is required to be 44pF. In the present example, this can be realized by 8 capacitances of each 352pF in series. In case of a resonance frequency of 298 MHz (corresponding to 7.0T), the total required capacitance of the resonance coil reduces to only 2pF, which can be realized in the present example by 8 x 16pF in series.

Further continuing this practical example, the Bi field is proportional with the current through the coil loop. As a consequence, since the induced voltage is thus proportional with the dimensions of the loop a current of IA results in an induced voltage of 57 volts at 1.5T, 112 volts at 3T and 262 volts at 7T. Thus, in order to avoid such high voltages at an individual capacitor, the many capacitors which are distributed in series over the coils geometry are dividing the high voltages onto the individual capacitors.

A possibility to limit the current is to place a 'high input impedance' preamplifier in series with the loop. Thus, due to the minimized current the inductive coupling of the other RF loops located nearby, for example in a coil array or in coil combinations, is minimized.

Fig. 10 is a schematic illustrating a nuclear magnetic resonance coil comprising a set of capacitances C, inductances L and resistances R. Further, a preamplifier 1000 is placed in parallel with the coil loop depicted in fig. 10. For optimal noise matching it is necessary to match the resonance loop frequency to the optimal input impedance of the preamplifier 1000. One way of doing this is using a 'resonance readout' by taking the voltage over one capacitor and dividing capacitance in such a way that the total equivalent series resistance in the coil loop when this is expressed as a parallel resistance is divided into the optimal impedance the preamplifier is tuned to for its optimal noise figure matching. In a practical example, the total inductance of the coil depicted in fig. 10 may be again assumed to be 14OnH and the total capacitance may correspond to 2pF (e.g. at 7T). Because, as already discussed with respect to fig. 9, the voltage may have been divided in 8 segments of 50mm length, this requires 8 capacitors equally distributed over the full length of 16pF. As a consequence, Cl in series with C2 must be 16pF. The required impedance can be obtained by choosing the right ratio in capacitance for Cl and C2.

Fig. 11a further considers additionally a series resistance typically caused by the patient loading of the coil. This resistance in series is indicated by Rs. In practice at a magnetic field of 7T, the Q factor in the present example can be calculated to Q = (ω x L) / Rs = 26. In other words, by means of the patient loading the Q factor of the coil drops. In fig. 1 Ib, this series resistance is translated into a parallel resistance Rp.

Mathematically, this can be done by Rp = Rs x Q² = 6800 Ω. Assuming an optimal impedance of the preamplifier of C_opt = 680 Ω, the impedance of the circuit depicted in fig. 1 Ib must be divided by a factor of 10. Such a division by 10 can be achieved by means of a 'capacitive Tap' Cl in series with C2. As already discussed with respect to fig. 9, at a magnetic field of 7T and the exemplary coil dimensions, the total capacitance of the circuit depicted in fig. 1 Ib needs to be 2pF. In fig. 10, this would accordingly correspond to a capacitance of the capacitor C2 of 16pF. For the capacitive Tap in fig. 1 Ib the ratio between Cl and C2 is chosen to be about 10:1. The capacitor with the highest capacitance gets the lowest voltage. Thus, Cl can be chosen to be 2OpF and C2 = 2pF. A realization of 2pF is rather difficult in discreet components, since the realization of such small capacitances also requires a consideration of the layout like for example the printed circuit board material used to carry the discreet components. However, in the present embodiment in case the base material of the printed circuit board is used as the dielectric material for usage as capacitor itself, only the layout and thus the electrical behavior of the printed circuit board itself needs to be considered.

A practical realization of such a 'capacitive Tap' is depicted in fig. 12. On the top of fig. 12, the coil loop segments 100 and 102 are sandwiching the dielectric material 202. By shifting the overlapping areas between the alternately arranged coil loop segments, the capacitances of Cl = 2OpF and C2 = 2pF can be realized. An alternative is depicted on the bottom of fig. 12, where the overlapping areas between the first coil loop segment 100 and the second coil loop segment 102 on the left hand side are different from the overlapping areas between the second coil loop segment 102 and the first coil loop segment 100 on the right hand side. It is further possible to combine a shifting of the overlapping regions in combination with varying widths of the coil loop segments 100 and 102, respectively. Thus, by for example accordingly adapting the layout design of a printed circuit board, capacitive taps for coil loops can be formed in a smart, simple and exact manner.

Fig. 13 depicts another embodiment in which a high input impedance preamplifier is used, however, in contrast to fig. 10 the preamplifier 1000 is put in series with the antenna loop. This allows for a preamplifier decoupling: by placing a high input impedance preamplifier in series with the antenna loop, the current that flows inside the antenna loop can be minimized and thus inductive coupling with other coils or coil elements being part of a coil array can be minimized. It has to be noted here, that limiting the induced current also minimizes the effect of inductive coupling with other kinds of electronic components nearby.

With an impedance matching network (PI matching network Cl, Ll and C2) it is possible to match the optimal impedance of the high input impedance preamplifier 1000. In practice, Cl can be part of the circuit which was built by means of the capacitance with overlapping areas. Ll can be a conductive path with a certain length and C2 can also be manufactured by means of the technique of overlapping areas.

Preferably, the matching circuit and the preamplifier must be close to the coil loop for the reason of an optimized signal-to-noise performance which is due to the fact, that losses before the first pre-amplification are avoided. Also, preferably the preamplifier is positioned close to the PI matching network before else the connection to the preamplifier will modify the impedance in an unwanted manner. By using a printed circuit board as a dielectric material layer which is sandwiched between conductive path thus forming a capacitor, discreet components in between the preamplifier and the coil loop can be avoided. REFERENCE NUMERALS

100 first coil loop segment

102 second coil loop segment

104 capacitor material

106 soldering point

108 active capacitor material

202 dielectric material layer

300 connection points

302 substrate

304 magnetic field direction

700 coil

702 coil

800 capacitor

802 capacitor

804 coil

806 coil

808 first conductive area

810 second conductive area

1000 preamplifier

Claims

CLAIMS:

1. A nuclear magnetic resonance coil comprising electrically conductive coil loop segments (100; 102), wherein the coil loop segments are alternately arranged on opposed sides of a dielectric material layer (202), wherein consecutively arranged coil loop segments (100; 102) are partially overlapping each other.

2. The coil according to claim 2, wherein the coil loop segments have the same segment length, wherein the consecutively arranged coil loop segments are overlapping each other in equal shares.

3. The coil according to claim 1, wherein the coil comprises a coil support with a top and a bottom side, wherein the coil loop segments are carried by the coil support on the top and bottom side, wherein the dielectric material layer (202) is formed by the coil support.

4. The coil according to any of the previous claims, wherein said coil comprises a primary magnetic field direction (304), wherein the coil loop segments have a planar structure, wherein the perpendicular of the planar structure is parallel to the primary magnetic field direction.

5. The coil according to any of the previous claims, wherein said coil comprises a primary magnetic field direction (304), wherein the coil loop segments have a planar structure, wherein the perpendicular of the planar structure is orthogonal to the primary magnetic field direction.

6. The coil according to any of the previous claims 3 to 5, wherein the coil is a printed circuit board.

7. The coil according to any of the previous claims, wherein the coil loop segments comprise a first, second and third segment, wherein the first and the third coil loop segments are arranged on a first side of the dielectric material layer (202) and wherein the second coil loop segment is arranged on a second side of the dielectric material layer (202), the second side being opposed to the first side, wherein the second coil loop segment is partially overlapping the first and the third coil loop segment.

8. A nuclear magnetic resonance coil array comprising a first coil (804) according to any of the previous claims and a second coil (806) according to any of the previous claims.

9. The coil array of claim 8, wherein the first coil (804) is spatially separated from the second coil (806), wherein the first (804) and the second (806) coil are decoupled from each other by a decoupling capacitance (800; 802), wherein the decoupling capacitance is formed by a first (808) a second (810) conductive area and a dielectric decoupling material layer (202), wherein the first conductive area is electrically connected to a coil segment of the first coil and wherein the second conductive area is electrically connected to a coil segment of the second coil, wherein the first conductive area is partially overlapping the second conductive area in an overlapping region, wherein the dielectric decoupling material layer is located in the overlapping region.

10. The coil array of claim 9 wherein the first conductive area (808) is carried by the coil support on the top side and wherein the second conductive area (810) is carried by the coil support on the bottom side, wherein the dielectric decoupling material (202) layer is formed by the coil support (202).

11. The coil array of claim 8, wherein the first and the second coil are overlapping for a mutual electromagnetically decoupling of the first and the second coil, wherein the first and the second coil are located in different layers of a multilayer printed circuit board.

12. A method of manufacturing a nuclear magnetic resonance coil comprising electrically conductive coil loop segments (100; 102), wherein the coil loop segments are alternately arranged on opposed sides of a dielectric material layer (202), wherein consecutively arranged coil loop segments are partially overlapping each other, the method comprising: providing the dielectric material layer (202), attaching the coil loop segments (100; 102) to the dielectric material layer (202).

13. A computer program product comprising computer executable instructions to perform any of the method steps as claimed in claim 12.