WO2010057564A1

WO2010057564A1 - Multi-channel housing device for mr imaging

Info

Publication number: WO2010057564A1
Application number: PCT/EP2009/007718
Authority: WO
Inventors: Jörn EWALD; Florian Meise; Torsten Hertz
Original assignee: Lmt Medical Systems Gmbh
Priority date: 2008-11-18
Filing date: 2009-10-28
Publication date: 2010-05-27
Also published as: DE202008015239U1

Abstract

The multi-channel housing device for magnetic resonance tomography (MRT) with several arrangements of partially overlapping rectangular coils (1-8), arranged on one or more coil carriers, characterised in comprising two domed arrangements of partly overlapping rectangular coils (1-8).

Description

Multi-channel recording device for MR imaging

The invention relates to a multi-channel recording device for magnetic resonance tomography (MRT) with a plurality of arrangements of partially overlapping rectangular coils, which are arranged on one or more coil carriers.

In general, such recording devices are referred to as MR coils. If one brings such a coil arrangement in the vicinity of the patient to be examined - e.g. through surface contact - this is also known as local coils. By the MR coils, the nuclear magnetic resonance signals are picked up by nuclear magnetic cores (primarily protons, i.e., nuclei of hydrogen) upon excitation thereof by radio frequency. The frequency of these signals depends on the magnetic field strength. By superimposing a strong magnetic field of typically 1.5 Tesla through one or more magnetic fields with gradients, the nuclear magnetic resonance signals can be measured at various points of the body and mosaically combined to form an overall image.

Examination of the body of newborns and infants on MRI places special demands on the local coil. Examination with the body module integrated in the MRI usually does not lead to sufficient image quality. This is mainly due to the excessive view ^¬ field, of too great a distance from the patient and the small Volume of the patient. The signal-to-noise ratio and the resolution of tissue details is too low. The same applies to the standard local coils, which are provided as accessories for the MRI. Although these coils cover a wide range of applications for the examination of adults and children, the suitability for neonates is very limited analogous to the reasons mentioned above.

The construction of a local coil as a multiple coil (array antenna, US Pat. No. 4,825,162) is known. This consists of an arrangement of juxtaposed partially overlapping individual coils. By means of the individual coils, the nuclear magnetic resonance signals of different body regions can be recorded substantially separated from one another, which optionally results in an improved overall image or a shortened examination time (parallel imaging). By combining a plurality of relatively small area MR coils into an array, the benefits of a large field of view and a high signal-to-noise ratio are brought into coincidence. The individual coils have a simple geometric and planar structure, for example as a loop or toroidal coil. The mutual influence of adjacent coils is greatly reduced, inter alia, by the partial overlap of adjacent coils.

The object of the invention is to provide a receiving device of the type mentioned, which is particularly suitable for infants and neonates (newborns). The solution according to the invention is that it has at least two arrangements of partially overlapping rectangular coils which are curved.

It is thereby possible a close concern to the body of the toddler, resulting in better signals. In an advantageous embodiment, at least one of the rectangular coils is flexible.

In particular, one of the arrangements of rectangular coils is attached to a coil support, which is designed as a slightly curved patient bed, while the other arrangement of rectangular coils is attached to a coil carrier, which can be adapted as a particular padded element to the body of the patient. The trained as a patient bed base takes into account the desired imaging of the spine from the neck area to at least down to the rump. The longitudinal axis of the patient couch corresponds to the z-axis and the orientation of the main magnetic field B ₀ in the MRI.

The flexible other coil carrier, the upper part, can be placed on the regions to be examined - the thorax or abdomen / pelvis of the patient. Due to the flexible structure which is flexible in the xy plane and padded as a cushion, the contour can be adapted to the patient in an advantageous manner and, if appropriate, be displaced in the z-direction into the desired field of vision.

Another advantageous embodiment is characterized in that the rectangular coils are arranged on a cylindrical coil carrier. Also in this case The rectangular coils enclose the child's body at a small distance.

With appropriate dimensioning, this embodiment is particularly suitable for examining the child's head.

The compact design of the coil carriers and the coils creates the possibility of using the arrangement in an MR-compatible incubator.

In an advantageous embodiment, more than one arrangement of rectangular coils are provided on at least one coil carrier, in particular on the couch formed.

In order to reduce the coupling capacity of the coupling of adjacent MR coils, it is advantageously provided that, in the region of the intersection of the tracks of the coils, they are substantially narrower than in the remaining part of the coils and intersect at 45 °. In an advantageous embodiment, adjacent coils are alternately provided on the front and the back of a printed circuit.

Advantageously, preamplifiers for the coils are each provided in the interior of the coils. The space-saving positioning of the preamplifier on the inner surface of the coil, in addition to a compact design of the coil has the advantage of a very short lead for the signal to be amplified per channel. Advantageously, preamplifiers of two adjacent coils are provided in pairs in the interior of each second coil, wherein the interior of each other second coil is transparent or air-permeable due to a corresponding recess of the coil carrier, for example, to facilitate breathing.

Advantageously, means are provided for detuning the resonant frequency of individual coils. It can thereby be achieved that the coil is not excited to resonance during the transmission phase of the MRI.

Advantageously, the coils have an overcurrent fuse, in particular a fuse. This fuse protects in case of error - e.g. at very high current increase in the conductor of the coils, caused by an (unintentionally) uncoordinated coil during the transmission phase of the MRI - the patient from burns.

Sheath wave barriers are advantageously provided by which signals are suppressed which are generated by eddy currents generated in the transmission phase in lead shields.

The invention will be described below with reference to an advantageous embodiment with reference to the accompanying drawings, for example. Show it:

FIG. 1a shows an arrangement according to the invention in the open position; FIG.

FIG. 1b shows the arrangement of FIG. 1b in the operating position; FIG. Fig. 2a shows the topology of a configured as a shell

Coil arrangement with N x n = l x 4 = 4 coils, thus four channels;

2b shows the topology of a coil arrangement configured as a lower part with N x n = 2 × 4 = 8 coils, that is to say eight channels;

3 shows a detailed representation of the overlap of the conductor tracks of the coils;

4 shows the combination of two amplifier boards to form a common double amplifier board using the coils rotated through 180 ° to each other;

5 shows an exemplary ring antenna with a downstream amplifier circuit;

6a is an illustration of a standing wave barrier;

FIG. 6b a representation of a standing wave barrier according to an alternative concept to FIG. 6a;

FIG. 7 shows a circuit implementation of the block diagram from FIG. 5; FIG.

FIG. 8 shows the frequency response of the RF blocking filter (7.1) from FIG. 7; FIG.

9 shows a topology according to the invention as an extension of FIG. 2a in the configured coil arrangement with N xn = 1 × 8 coils; So with 8 channels as the lateral surface of a cylinder; and Fig. 10 shows the topology of Fig. 9, with the schematic

Arrangement of the four amplifier boards and four cut coil inner surfaces as a viewing window.

Fig. Ia and Ib show the basic concept of the invention. The multichannel body coil or its coil carrier is subdivided into an upper part (Ia.1) which is placed on the regions to be examined - the thorax or abdomen / pelvis of the patient (Ib.1). Due to the flexible construction, which is flexible in the xy plane and padded as a cushion, the contour can be adapted to the patient in an advantageous manner and optionally displaced in the z-direction into the desired field of vision.

At the same time, the lower part (Ia.2) is shaped as a patient couch surface slightly curved in the xy plane, which favors use in the MR incubator. The geometry of the base takes into account the desired imaging of the spine from the neck area down to at least the rump. It should be noted that the z-axis also represents the orientation of the main magnetic field B ₀ in the MRI.

Thus one clearly sees the carriers of the two partial coils of the multi-channel body coil, namely the flexible upper part (Ia.1) and the rigid lower part (Ia.2), which at the same time represents the patient lying surface. In FIGS. 1a and 1b, a coordinate system is also defined for further reference. FIGS. 2a and 2b show the array geometries of the coils from which the two coil arrangements are constructed. The brightly drawn coils are printed on the upper side of the board, corresponding to the dark drawn on the lower side of the board on which they are arranged. It is characteristic that the same geometry is used for all coil arrangements. This means for the application in neonates a width d = 50mm in the x-direction and a length l = 125mm in the z-direction. The width of the tracks is predominantly 5mm. For the flexible upper part, the structure according to FIG. 2a is shown. The coil arrangement consists of four coils 1 to 4 overlapping in the x-direction.

Each coil constitutes a conductor loop with a self-inductance L _A predetermined by its geometry.

If two such conductor loops are brought into spatial proximity, these two loops will couple each other via the counter-inductance M. Furthermore, the overlapping strip conductors of the front and back of the printed circuit act as coupling capacitance C due to the opposite cut surfaces. Both types of coupling are undesirable, since the image reconstruction in the MRI system (image computer) emanates from decoupled individual signals of the ring antennas. Any electromagnetic coupling of adjacent loop antennas would significantly increase noise in the image area (US 4,825,162). The mutual inductance M can be minimized by overlapping the coils. Δ for a given width D, and the overlap is (for the optimum M ^ O), the ratio δ / D = 0.14 known (for square Lei ^¬ terschleifen). For the geometry defined above δ / D = 0.21 was determined experimentally. The coupling capacitance C is drastically reduced by a local rejuvenation of the trace (here: from 5mm to 1mm) in the overlap area. In addition, by 90 ° alignment of the interconnects, the intersecting surface segments can be reduced to ultimately A = lmm ² . This has a very advantageous effect on minimizing the capacitive coupling because of C - A.

Fig. 3 illustrates these relationships in detail, namely the details in the areas of the overlap δ and overlap area A of two adjacent coils. The light-colored conductor segments are printed on the upper side of the circuit board, corresponding to the shaded areas on the lower side of the circuit board.

2b shows the structure of the lower, rigid arrangement of coils 5 to 12. Here, the sub-array of Fig. 2a is used twice - each offset in the z-direction. Also in this direction, the decoupling of pairs opposite coils is achieved by the overlap δ.

FIG. 4 shows how the signal respectively received by the coils (4.1) and (4.2) via the symmetrical connection (4.4) resp. (4.5) is fed into the common double preamplifier board (4.3). This board has a point-symmetrical structure with regard to the components as the preamplifier (4.6) resp. (4.7) and standing wave barriers (4.8) resp. (4.9). The supply lines (DC and RX from FIG. 5) are connected to the two short sides of the board (4.3). The pairwise combination of the preamplifier boards is advantageous because the now free inner surface of the loop antenna (4.2) can be cut out and thus a light and air-permeable window to the patient is created.

In Fig. 5, the structure of a channel of the preamplifier board is shown on the level of function blocks. The final received signal to the MRI is RX; With the control signal DC coming from the MRT, the coil can be deliberately detuned and thus switched off. The coil (5.1) is supplemented by a tuning circuit (5.2), which is realized in the simplest case with a capacitor C _A. This forms in series with the two series-connected capacitors in the balancing circuit (5.3) C _s, the capacitive component to the inductively embossed part L _{A of} the conductor loop of the coil. The resonance frequency can thus be represented as follows:

In practice, the C _{s are} fixed and set the entire coil by means of C _A by adjustment or targeted value selection to the resonant frequency.

In the event of a fault, the fuse (5.8) protects the patient from burns, for example when there is a very strong current increase in the conductor of the coil caused by an (unintentionally) unsanctioned antenna during the transmission phase of the MRI. The balanced signal applied to the coil is asymmetrically decoupled by means of the balancing circuit (5.3), ie related to ground.

This signal is fed into the preamplifier (5.6), but via a decoupling circuit (5.5) located in its input path. In the simplest case, this is a capacitive component -j / ωC with which the coupling-active component + jcύM ⁽ (caused by the remaining interaction M 'between adjacent ring antennas) is completely compensated.

In order to achieve this, the lowest possible input impedance Zi of the preamplifier (5.6) is necessary. With this, the output of the decoupling circuit would ideally be grounded. This situation is also described in US 4,825,162.

The preamplifier itself is a modular component with a fixed voltage gain v. This is dimensioned so that the signal received by the antenna on the from the

MRI manufacturer specified level is raised, for example, v = 3OdB is a typical value.

An important feature of the preamplifier is a very low noise level; because this parameter directly affects the picture quality. The output impedance is Z ₀ = 50 ohms in most cases; As a rule, the preamplifier is also supplied with a DC voltage superimposed on this output.

The structure of the standing wave barrier (5.7) is shown in Fig. 6a. The signal coming from the preamp output is shielded signal cable (6a.1) is wound on a toroidal winding body (6a.2) made of Teflon (in real about 18 turns). Thus, the shielding is effective as inductance L _MWS . At the entry and exit points of the cable, the screen is bridged by a tunable capacitor C _MWS (6a.4) and hereby a blocking filter for f _{0 is} realized via this parallel resonant circuit.

This provides effective suppression of on-induced, damaging mantle wave currents. The signal is thus conducted via the inner conductor (6a.4) of the coaxial cable (6a.1) wound on the toroidal body (6a.2). At the two end points, the braided screen is bridged with an adjustable capacitor (6a.3). Hereby, the blocking frequency is adjusted to f o and thus the sheath current is effectively suppressed, which is caused by eddy currents generated in the shield during the transmission phase by the radiated high frequency.

An alternative construction of the standing wave barrier is shown in FIG. 6b. Here, the coaxial cable (6b.1) is wound on a cylindrical winding body (6b.2) instead of a toroidal body. The winding body is hollow on the inside and takes up a copper core (6b.5) whose immersion depth is e.g. is adjustable via a thread.

Since copper has diamagnetic properties (χ <0), the more the copper core dips into the coil body, the lower the effective inductance L _MWS . This, in turn, causes an increase in the resonance frequency (equal to blocking frequency). frequency of the standing wave barrier) after itself. The capacitor (6b.3) can therefore be provided as a fixed value CMWS; the resulting resonance frequency f ₀ is again calculated according to the above formula.

The detuning circuit (5.4) is explained in more detail by circuit diagram Fig. 7. A DC control current coming from the MRT (approx. 100mA) is conducted via an HF cut-off filter (7.1) via the current path formed by the PIN diode (7.2) and the detuning inductance (7.3) to ground. In the forward direction, the PIN diode is virtually a closed switch due to its very low differential internal resistance, which makes the detuning inductance (7.3) act in the resonant circuit of the coil. Thus, the resonance frequency is shifted from f ₀ to £ χ. Since the MRI system registers only resonances and signals near f ₀ , the respective tuned channel is invisible to MRT. If the control current from the MRT is set to 0, then the PIN diode is pulled out of the resonant circuit with high resistance (open switch) and the detuning inductance. Thus again, it is effective and the channel is visible again for the MRI. The Verstimminduktivitat (7.3) can for example be designed so that it corresponds to the inductance L _{A of} the coil. In the case of detuning, the total inductance is thus 2L _A , which implies:

The detuning frequency fi is thus far enough away from the actual MRT resonance frequency f ₀ . The RF rejection filter (7.1) prevents short circuiting of the high frequency signals from the coil's point of view through the DC source of the MRI; From the point of view of the MRI, (7.1) prevents the coupling of high-frequency signals; its frequency response is shown in Fig. 8.

The balancing circuit (7.4) can be seen as a capacitive voltage divider with the two capacitors C _s = 18OpF, with a grounded center tap. The lower tap thus represents the now asymmetric (ground related) output of the otherwise symmetric coil.

In the components referenced in FIGS. 5 and 7, (5.4) corresponds to the combined components (7.1) + (7.2) + (7.3); (5.2) corresponds to (7.6); (5.3) corresponds to (7.4); (5.5) corresponds to (7.7); (5.6) corresponds to (7.8) and (5.7) corresponds to (7.9).

9 shows a coil configuration which is based on the arrangement of FIG. 2a. The arrangement extended in the x-direction to 8 individual coils is rolled up to form a cylindrical outer surface, so that the first and last individual coil again overlap with δ, thus resulting in a rotationally symmetrical structure.

In the center of such a hollow cylinder can be advantageously, e.g. placing the head of a newborn, which combines the benefits of high image resolution, high signal-to-noise ratio, and the potential for parallel imaging.

This is made possible by the use of flexible conductor carrier material (eg Pyralux film), which The sides are covered with copper and can be processed and through-contacted.

The exemplary cylindrical dimensions of the multi-channel coil for neonate head examinations have a diameter of D = 175 mm; for the height h = 120 mm. The width of a single coil was set at d = 86.5 mm; the optimal overlap was determined empirically with δ = 16.5 mm.

10 shows the advantageous arrangement a) of the amplifier boards (10.1, four boards according to FIG. 4) in the region of the individual coils 2, 3, 6 and 7 and b) of the viewing windows (10.2) in the region of the individual coils 1, 4, 5 and 8th.

The proximity of the amplifier boards to the individual coils has an advantageous effect on the signal quality.

The viewing windows advantageously allow an immediate observation of that located in the head coil

Patients, in addition to the already open end faces of the coil assembly.

Claims

protection claims

1. A multi-channel recording device for magnetic resonance imaging (MRI) with a plurality of arrangements of partially overlapping rectangular coils (1-12), which are arranged on one or more coil carriers (Ia.1, Ia.2), characterized in that they at least two curved arrangements partially overlapping rectangular coils (1-12).

2. Recording device according to claim 1, characterized in that at least one of the rectangular coils (1-

12) is arranged flexibly.

3. Aufnähmeeinrichtung according to any one of claims 1 or 2, characterized in that the rectangular coils (1 to 12) are arranged on a cylindrical coil carrier.

4. Recording device according to claim 1, characterized in that one of the arrangements of rectangular coils (5-12) is attached to a coil carrier (Ia.2), which is designed as a slightly arched patient bed, while the other arrangement of rectangular coils (1- 4) is attached to a coil carrier (Ia.1), which is adaptable as a particular padded element to the body of the patient.

5. Recording device according to claim 1 or 2, characterized in that on at least one coil carrier (Ia.2), in particular on which is designed as a couch, more than an array of rectangular coils (1-12) is provided.

6. Recording device according to one of claims 1 to 5, characterized in that in the region of intersections of the conductor tracks of the coils (1-12), these are formed substantially narrower than in the remaining part of the coils and intersect at 90 °.

7. Recording device according to one of claims 1 to 6, characterized in that adjacent coils (1- 12) are alternately provided on the front and the back of a printed circuit.

8. Recording device according to one of claims 1 to 7, characterized in that preamplifiers (4.6, 4.7, 5.6) are provided for the coils in each case in the interior of the coils (1-12).

9. Recording device according to one of claims 1 to 8, characterized in that preamplifiers (4.6, 4.7) of two adjacent coils (1-12) are provided in pairs in the interior of each second coil (1-12), wherein the interior of each other second coil (1-12) is transparent or permeable to air due to a corresponding recess of the bobbin.

10. Recording device according to one of Ansprüchel 1 to 9, characterized in that means for

Detune (5.4) of the resonant frequency of individual coils (1-12) are provided.

11.Aufnahmeeinrichtung according to any one of claims 1 to

10, characterized in that the coils (1-12) have an overcurrent fuse, in particular a fuse (5.8).

12.Aufnahmeeinrichtung according to any one of claims 1 to

11, characterized in that standing wave barriers (5.7) are provided, are suppressed by the signals which are generated by generated in the transmission phase eddy currents in Zuleitungsabschirmungen.

13.Aufnahmeeinrichtung according to any one of claims 1 to

12, characterized in that adjacent coils (1-12) for minimizing coupling inductances overlap by a defined distance δ.

14.Aufnahmeeinrichtung according to any one of claims 1 to

13, characterized in that the conductor tracks of adjacent coils (1-12) for minimizing

Coupling capacities overlap only by a minimum area A in the x, z plane.

15.Aufnahmeseinrichtungen according to any one of claims 1 to 14, characterized in that each preamplifier (4.6, 4.7, 5.6) of each coil (1-12) the substructures detuning circuit (5.6), decoupling circuit (5.5), preamplifier (4.6, 4.7, 5.6 ) and standing wave barrier (4.8, 4.9, 5.7).

16.Aufnahmeeinrichtung according to any one of claims 1 to 15, characterized in that in the conductor path of each coil (1-12) the substructures fuse (5.8), sym- ming circuit (7.4) and tuning circuit (5.2) are inserted.

17.Aufnähmeeinrichtung according to any one of claims 1 to 16, characterized in that for all channels of a coil (1-12), the shielded coaxial line of the EmpfangssignaIs and the control for the detuning circuit (5.6) are combined in a common shielded connection cable.