Low-field magnetic resonance imaging system

Abstract

The invention relates to a magnetic resonance imaging system comprising at least one radiofrequency gradiometer (GR), which comprises two conductive coils (BCA, BCB), referred to as capture coils, having a common sensitivity axis (z) and spaced apart along the axis, wherein each of the coils is connected in series to another conductive coil (BEA, BEB), referred to as the input coil, to form a flux concentrator (CF), and wherein each input coil is inductively coupled to a SQUID magnetometer (MSA, MSB); characterised in that the gradiometer also comprises an electronic processing circuit (CET) configured to linearly combine output signals (VSA, VSB) from the SQUID magnetometers in order to obtain a signal (VS) representative of a difference between radiofrequency magnetic field values (Bz) corresponding to the two capture coils.

Description

DESCRIPTION Title of the invention: Low magnetic field magnetic resonance imaging apparatus

[0001] The invention lies in the field of magnetic resonance imaging (MRI) and more particularly of very low magnetic field MRI.

[0002] We speak of very low magnetic field MRI when the stationary magnetic field of polarization of the nuclear spins has an intensity between approximately 50 pT and 100 mT, corresponding to Larmor frequencies of the order of kHz up to MHz for protons). Under these conditions, the amplitude of the magnetic field to be detected is typically less than 1 pT.

[0003] It is known from WO 2022 / 117969 to use, for the detection of such a weak MRI signal, a system based on SQUIDs (from the English "Superconducting Quantum Interference Device"), and more specifically low-temperature direct current SQUIDs (dc-SQUIDs). More information on SQUIDs and their use is provided by (Fagaly) and (Clarke, Braginski).

[0004] As illustrated in [Fig. 1], a DCS direct current SQUID consists of a superconducting loop BS interrupted by two Josephson junctions JJ1, JJ2. When a direct current I flows in the loop, the potential difference V across the Josephson junctions is a sinusoidal function of the magnetic field flux > sacross the loop. Typically, a bias coil BP is used to set an operating point around which the voltage varies approximately linearly with the flux. Advantageously, a feedback loop - called a flux-locking loop (FLL) - controls the current i p circulating in the polarization coil so as to keep the voltage V constant s supplied as input to the FLL loop after being amplified by a low noise amplifier LNA. In this case, the current i p - or another signal generated by the FLL loop and proportional to i p - can be taken as an output signal providing a measurement of the flow 4> s Alternatively, if the SQUID has sufficient dynamics, the FLL loop may have a bandwidth lower than the frequency of the useful signal; in this case, the output signal is constituted by a high frequency voltage variation across the SQUID terminals, V S HF .

[0005] Also, the assembly consisting of the SQUID DCS, the low noise amplifier LNA, the bias coil BP and the flux-locked loop FLL form an extremely sensitive MS magnetometer.

[0006] It is known to produce SQUIDs both with "conventional" superconductors, with low critical temperature, and with superconductors with high critical temperature. "Low critical temperature" means a critical temperature less than or equal to 20 K. In the aforementioned document WO 2022 / 117969, the use of SQUIDs with low critical temperature - maintained at a temperature well below the transition temperature, for example of the order of 4 K - is recommended due to the lower noise level, the greater ease of manufacture and the better figure of merit R n L c (where R n is the resistance of the junction in the normal state and l cits critical current) of these devices compared to high critical temperature SQUIDs. Furthermore, it is known that low-temperature superconducting materials can exhibit superior mechanical, shaping and stability properties than high-temperature superconductors.

[0007] The BS superconducting loop has a small surface area (a few pm 2 or tens of pm 2), it therefore intercepts a weak magnetic flux. Furthermore, as it must be kept at cryogenic temperature, it cannot always be optimally positioned to detect a magnetic field, for example near the body of a patient undergoing an MRI examination. For these reasons, it is known to use what is commonly called a CF flux concentrator. The CF flux concentrator comprises a first conductive coil BC, called the pickup coil, connected in series to a second conductive coil BE, called the input coil, the latter generally being made using planar technology and being inductively coupled to the superconducting loop BS of the SQUID.The mutual induction coefficient between the input coil BE and the superconducting loop BS is designated by MES- The pickup coil BC, which constitutes, in fact, an antenna, has an effective surface - product between the number of turns and the geometric surface of each turn - much greater than that of the superconducting loop, it is therefore crossed by a greater flux - flux which is transferred to the loop. superconducting SQUID through the input coil. Furthermore, the pickup loop does not have to be kept at cryogenic temperature and can therefore be positioned more freely.

[0008] The pickup coil BA is not necessarily a simple planar coil, as illustrated in [Fig. 1]. As explained in the aforementioned document WO 2022 / 117969, it can also be, for example, a volume coil, such as a saddle coil or a pair of Helmholtz coils. It can also be a gradiometric antenna. [Fig. 2A] shows a first-order gradiometric coil BG1 , which consists of two windings BG1 ', BG1 ” connected in series and wound in opposite directions and arranged. The windings BG1 ', BG1 ” are planar, located in planes perpendicular to a z-axis and spaced a distance d (so-called "baseline") along said axis. It is easy to understand that the current induced in the gradiometric coil BG1 depends on the variation of the component B z of the magnetic field on the baseline d, which - to a first approximation - is proportional to (3B zÆ)z)-d. The intensity of a magnetic field decreases as the cube of its distance from the source, but that of the magnetic field gradient decreases as the fourth power of this distance. Also, replacing a simple coil with a gradiometric coil makes it possible to attenuate the impact of distant noise sources. [Fig. 2B] represents a second-order gradiometric coil BG2, obtained by replacing the windings BG1 ' and BG1 ” with first-order gradiometric coils. It is easy to understand that the current induced in the gradiometric coil BG1 depends, as a first approximation, on the second derivative of B z relative to z: 3 2 B z / 3z 2 . Similarly, it is possible to produce higher order gradiometric coils, the rejection of noise from distant sources increasing with the said order.

[0009] [Fig. 2C] represents a so-called "saddle coil" BSC, used in MRI due to its field homogeneity properties. [Fig. 2D] represents the gradiometric (first-order) version of a saddle coil, BSCG, consisting of two saddle coil sub-antennas: a first, internal antenna, having - in the example shown - two turns of wire, and a second, external and larger antenna having - still in the example shown - a single turn of wire. The difference in the number of turns of wire allows the two sub-antennas to have the same inductance despite their size different. The windings are wound so that the current flows in opposite directions in the two sub-antennas.

[0010] A disadvantage of these gradiometric setups is that the antennas use long lengths of wire due to the succession of coils compensating each other. The capture antennas manufactured in this way generally have a relatively high inductance, which can pose a problem for the detection sensitivity with SQUID. Indeed, for quantum interference to be observed, the self-inductance L s of a SQUID must satisfy the condition

[0012] Or <j>where o is the flux quantum, T the absolute temperature and k B Boltzmann's constant. Typically, L s is therefore of the order of 1 pH.

[0013] If we call <TBC le flux à travers la bobine de captation BC et L B its inductance, the current i B c which runs through said coil and the input coil BE is worth

[0015] The flux <Î>SA, through the superconducting loop of the SQUID is worth

[0016] <P SA = M ESiBC (3)

[0017] With M ES = k jL E L s ( )

[0018] where L E is the inductance of the input coil BE and k is a dimensionless factor. We therefore find

[0020] It follows from equation (5) that, for a flow <ï> BC through the given pickup coil, the flux <T SA seen by the SQUID is maximal for L E =L BC . Now, to maximize 4> BC it would be desirable to take L BC large, while L E must be of the same order of magnitude as L s to maximize coupling (given the constraints inherent in its planar construction, the BE coil must have a geometry close to that of the SQUID to be able to couple inductively to it effectively). The low value of the SQUID inductance therefore imposes a limit maximum to that of the input coil, and therefore to that of the pickup coil. It follows that l_ B c cannot exceed a few pH or tens of pH, which is very restrictive in particular for the production of volume coils in an MRI scanner which, due to their dimensions, tend to have high inductance values. More particularly, it is preferable that the ratio L E / LBC should be between 0.4 and 2.5 so that the sensitivity is at least 90% of its maximum value.

[0021] The use of a gradiometric coil exacerbates this problem. Indeed, as shown above, such a coil has several windings, only a part of which is dedicated to capturing the signal of interest, the rest being dedicated to noise suppression. As a result, a large part of the inductance of the pickup coil - the value of which is limited by the adaptation constraints explained in the previous paragraph - is not used to collect the signal of interest. This problem is even more serious for higher-order gradiometers.

[0022] In the aforementioned document WO 2022 / 1 17969, a transformer is interposed between the pickup coil and the input coil. This transformer makes it possible to increase the impedance L A while respecting the adaptation constraints. However, it presents significant losses, especially when the transformation factor is high, which limits the interest of this solution. In addition, transformers having to work at low temperatures can hardly use ferromagnetic cores.

[0023] The invention aims to overcome at least in part the aforementioned drawbacks of the prior art.

[0024] According to one aspect of the invention, this object is achieved by using an electronic processing circuit configured to linearly combine the output signals of two independent SQUID magnetometers to obtain a signal representative of a difference between magnetic field values ​​corresponding to the pickup coils of said magnetometers. In this way, the entire inductance of each pickup coil can be used to acquire a useful signal. Alternatively, each of the two magnetometers can have a gradiometric antenna of order N (N > 1), the electronic processing circuit making it possible to obtain a gradiometer of order N+1.

[0025] A similar setup was disclosed in (Matlashov) in the case of a device based on high critical temperature SQUIDs and with superconducting acquisition coils made in planar technology. In this case, indeed, it is difficult for technological reasons to realize gradiometric setups with a sufficiently high baseline.

[0026] Similarly, JP 2010 151508 discloses the gradiometer comprises two coaxial pickup coils, spaced in an axial direction, inductively coupled to respective SQUIDs whose output signals are subtracted by an electronic circuit. This gradiometer, however, is not designed for radio frequency application, let alone for MRI.

[0027] JP H06 138197 discloses a gradiometer using two SQUIDs associated with respective pickup coils, a signal representative of the difference between the magnetic fluxes through the two pickup coils being obtained by coupling the flux-locked loops. Again, the device is not designed for radiofrequency application, and even less so for MRI.

[0028] US 5,355,085 discloses a multi-channel magnetometer comprising a plurality of SQUID sensors connected to a multiplexer.

[0029] An object of the invention is therefore a magnetic resonance imaging apparatus (MRIA) comprising at least one radiofrequency gradiometer comprising two conductive coils, called pickup coils, having a common sensitivity axis and spaced along said axis, each of said coils being connected in series to another conductive coil, called input coil, to produce a flux concentrator, each said input coil being inductively coupled to a SQUID magnetometer; characterized in that the gradiometer also comprises an electronic processing circuit configured to linearly combine output signals from said SQUID magnetometers to obtain a signal representative of a difference between radiofrequency magnetic field values ​​corresponding to the two pickup coils.

[0030] According to particular embodiments:

[0031] - Said pickup coils can be configured to operate in a frequency range between 1 kHz and 100 MHz.

[0032] - Said SQUID magnetometers may comprise direct current SQUIDs.

[0033] - Said SQUID magnetometers may comprise low critical temperature SQUIDs.

[0034] - Said electronic processing circuit may be an analog circuit comprising an adder or subtracter amplifier for adding or subtracting said output signals. More particularly, each said SQUID magnetometer may comprise a low-noise amplifier for generating the corresponding output signal by amplifying a voltage drop across a respective SQUID.

[0035] - Each said magnetometer may include a flux-locked loop.

[0036] - Each said pickup coil can have an inductance greater than or equal to 10 pH.

[0037] - Said capture coils can form volume antennas.

[0038] - The device may also include a coil for generating a stationary magnetic field for polarizing nuclear spins, with an amplitude between 10 pT and 100 mT.

[0039] Other characteristics, details and advantages of the invention will emerge from reading the description given with reference to the appended drawings given by way of example and which represent, respectively:

[0040] [Fig.1], described above, a SQUID magnetic field detection system known from the prior art.

[0041] [Fig. 2A], [Fig. 2B], [Fig. 2C] and [Fig. 2D], described above, different types of antennas that can be used with the detection system of [Fig. 1] or for implementing the invention;

[0042] [Fig. 3], a radiofrequency gradiometer according to a first embodiment of the invention;

[0043] [Fig. 4], a radiofrequency gradiometer according to a second embodiment of the invention; and

[0044] [Fig. 5], a nuclear magnetic resonance imaging device (scanner) using at least one radiofrequency gradiometer according to one embodiment of the invention.

[0045] The apparatus of [Fig. 3] comprises two SQUID magnetometers analogous to that of [Fig. 1], MS A , MS B . Each of these magnetometers comprises a direct current SQUID DCSA, DCSB inductively coupled to an input coil BE A , BE B as well as a BP polarization coil A , BP B fixing its operating point and driven by a flux-locked loop not shown. The input coil is connected in series to a pickup coil BCA, BCB- These two pickup loops are arranged like the two windings of the gradiometric antenna in [Fig. 2A]: they are planar, arranged in planes perpendicular to the z direction (“sensitivity axis”) and therefore sensitive to the B component z of an external magnetic field; they are furthermore spaced by a distance d in said z direction. If z0 is the coordinate along the z axis of the midpoint of the two coils, the coil BC A is located at position z0+d / 2 and the coil BC B at position z0+d / 2.

[0046] The voltages across the two SQUID V A , V B are amplified by low noise amplifiers LNA A , LNA B . The directions of the windings of the BC pickup coils A , BC B and BE input coils A , BE B are chosen so that a magnetic flux through the pickup coils induces voltages V A , V B of opposite sign. For example, pickup coils may have windings wound in opposite directions and input coils in the same direction, or vice versa.

[0047] V exits SA , V SB low noise amplifiers are provided as input to an electronic processing circuit CET which, in this embodiment, is a simple voltage adder circuit (more precisely a non-inverting adder) comprising an operational amplifier OA and four resistors Ri, R2, R3, R4. If RI=R2, the output voltage V s of the operational amplifier OA is proportional to VA+V B and therefore the difference in magnetic fluxes across the two pickup coils:

[0049] The system therefore constitutes a first-order gradiometer, identified by the reference GR. Unlike the case of the gradiometer in [Fig. 1], however, the entire inductance of each of the two pickup coils is used to acquire a useful signal; this architecture therefore makes it possible to alleviate the constraints on the value of this inductance.

[0050] In more detail, if L B C=LE (we consider that the two pickup coils BC A , BC B have the same inductance L B c, the two BE input coils A , BE B the same inductance L E and the two SQUID DCS A , DCS B the same inductance Ls) equation (5) becomes

[0052] Where <Î>BC is worth depending on whether we consider one or the other of the two pickup coils.

[0053] The voltage V A at the DCS terminals A is therefore worth

[0055] Where K s is the flux-voltage transfer coefficient of the SQUID (assumed identical for DCS A and DCSB). Similarly, the voltage V B at the terminals of DCSB is worth

[0057] By introducing a parameter A which takes into account the LNA amplifiers A , LNA B (assumed identical), from the effect of the flux-locked loop and the gain of the voltage adder circuit - equal to 1 +R3 / R4 - we obtain:

[0059] In equation (10), the last equality was obtained by considering that the two pickup coils have the same effective surface area SBC (in the simple case of a planar coil, this is the surface area projected onto the xy plane perpendicular to the z axis and multiplied by the number of windings) and that therefore <Ï>=B Z SBC-

[0060] In the embodiment of [Fig. 4], the directions of the windings of the pickup coils BC A , BC B and BE input coils A , BE B are chosen so that a magnetic flux through the pickup coils induces voltages V A , V B of the same sign. These voltages are supplied to the differential inputs of a low-noise amplifier LNA'. As in the previous case, the output voltage V s is proportional to

[0062] As in the case of [Fig. 3], the system therefore constitutes a first-order gradiometer, identified by the reference GR'.

[0063] [Fig. 5] represents very schematically a device (or "scanner") for magnetic resonance imaging SMRI. Such a device comprises a BCM coil for generating a static and uniform magnetic field B o oriented in a so-called longitudinal direction, which makes it possible to orient (polarize) in said direction the nuclear spins of the atoms of a sample - for example a part of a human body, or the entire body. Coils, called gradient coils (not shown) apply static and non-uniform magnetic fields, necessary to obtain spatial resolution. Radiofrequency coils, not shown, make it possible to apply excitation pulses to said nuclear spins, following which the nuclear spins de-excite by emitting radiofrequency radiation. One or more GR gradiometers according to the invention, equipped for example with volume coils surrounding the sample, make it possible to detect this radiation, their capture coils serving as receiving antennas.

[0064] The invention has been described with reference to certain embodiments, but several variants are possible. In particular:

[0065] - Other types of pickup coils can be used, such as volume coils. For example, the pickup coils can be the sub-antennas of a gradiometric saddle volume antenna, of the type in [Fig. 2D]. As mentioned above, these coils can also be N-order gradiometric coils (N > 1), in which case the processing electronics synthesize an N+1 order gradiometer. The inductance values ​​mentioned are given only as examples.

[0066] - Although it is generally preferable for both antennas to have the same inductance, it is also possible for them to have different inductances. Similarly, the SQUIDs and input coils of the two elementary magnetometers may not be identical. The resulting differences in output signal intensity must then be compensated for by the processing electronics. In the case of the embodiment in [Fig. 3] this is achieved by choosing different values ​​for resistors R1 and R2. Similarly, it is possible to compensate for any difference in effective area.

[0067] - The electronic processing circuit can add or subtract the signals it receives at the input (weighted where appropriate to take into account any difference between the two elementary magnetometers), depending on the winding directions of the pickup coils and the input coils.

[0068] - The processing electronic circuit may have a different topology than those described. For example, the processing may be partially or completely digital.

[0069] - Although the use of low critical temperature SQUIDs is preferred, the use of high critical temperature SQUIDs is not excluded.

[0070] - In some cases, the flux-locked loop can be omitted, the polarization of the SQUIDs being fixed. Conversely, the output signal of each elementary magnetometer can be provided by the flux-locked loop instead of being taken from the terminals of the SQUIDs.

[0071] - The invention can be implemented in other frequency ranges. In addition, the magnetic field polarizing the nuclear spins can be outside the intensity range indicated as preferential.

[0072] References

[0073] (Clarke, Braginski) John Clarke, Alex I. Braginski (editors) The SQUID Handbook, Vol. I (2004)

[0074] (Fagaly) R. L. Fagaly « Superconducting quantum interference device instruments and applications » Review of Scientific Instruments 77, 101101 (2006)

[0075] (Matlashov) A. Matlashov et al. « Electronic Gradiometer using HT C SQU IDs with Fast Feedbeck Electronics » IEEE Transactions on Applied Superconductivity, Vol. 11 , No. 1 , mars 2001.< / j>

Claims

CLAIMS 1. Magnetic resonance imaging device (MRID) comprising at least one radiofrequency gradiometer (GR, GR') comprising two conductive coils (BCA, BCB), called capture coils, having a common sensitivity axis (z) and spaced along said axis, each of said coils being connected in series to another conductive coil (BE A , BE B ), called input coil, to produce a flux concentrator (CF), each said input coil being inductively coupled to a SQUID magnetometer (MS A , MSB); characterized in that the gradiometer also comprises an electronic processing circuit (CET) configured to linearly combine output signals (VS A , VS B ) of said SQUID magnetometers to obtain a signal (V s ) representative of a difference between magnetic field values ​​(B z) at radio frequency in correspondence of the two capture coils.

2. Magnetic resonance imaging apparatus according to claim 1 wherein said pickup coils (BC A , BC B ) are configured to operate in a frequency range between 1 kHz and 100 MHz.

3. Magnetic resonance imaging apparatus according to one of the preceding claims wherein said SQUID magnetometers (BC A , BC B ) include direct current SQUIDs (DCS A , DCS B ).

4. Magnetic resonance imaging apparatus according to one of the preceding claims wherein said SQUID magnetometers (BC A , BC B ) include low critical temperature SQUIDs (DCS A , DCS B ).

5. Magnetic resonance imaging apparatus according to one of the preceding claims, in which said electronic processing circuit (CET) is an analog circuit comprising an adder or subtractor amplifier (AO, LNA') for adding or subtracting said output signals.

6. A magnetic resonance imaging apparatus according to claim 5 wherein each said SQUID magnetometer comprises a low noise amplifier (LNA A , LNA B ) to generate the corresponding output signal by amplifying a voltage drop (V A , V B ) at the terminals of a SQUID (DCS A , DCS B ) respective.

7. Magnetic resonance imaging apparatus according to one of the preceding claims wherein each said magnetometer (MS A , MS B ) includes a flux-locked loop (FLL).

8. Magnetic resonance imaging apparatus according to one of the preceding claims in which each said capture coil (BC A , BC B ) has an inductance greater than or equal to 10 pH.

9. Magnetic resonance imaging (MRI) apparatus according to one of the preceding claims in which said capture coils (BC A , BC B ) form volume antennas.

10. Magnetic resonance imaging device (MRID) according to one of the preceding claims comprising a coil (BCM) for generating a magnetic field (B o ) stationary polarization of nuclear spins, with amplitude between 10 pT and 100 mT.

Images