SQUID radio frequency gradiometer

Abstract

The invention relates to a radio frequency gradiometer (GR) comprising two conductive coils (BCA, BCB), called sensing coils, having a common sensitivity axis (z) and spaced along said axis, each of said coils being connected in series to another conductive coil (BEA, BEB), called input coil, to make a flux concentrator, each said input coil being inductively coupled to a SQUID magnetometer (MSA, MSB); characterized in that the gradiometer also comprises an electronic processing circuit (ETC) configured to linearly combine output signals (VSA, VSB) from said SQUID magnetometers to obtain a signal (VS) representative of a difference between radio frequency magnetic field values ​​(Bz) corresponding to the two sensing coils. The invention also relates to a magnetic resonance imaging device comprising at least one such radio frequency gradiometer, used as a receiving antenna.Figure for the summary: Fig. 3.

Description

Title of the invention: SQUID radiofrequency gradiometer

[0001] The invention relates to the field of ultrasensitive magnetic field sensors operating in the radio frequency range. It is primarily applicable to the field of magnetic resonance imaging (MRI) and more particularly to very low magnetic field MRI.

[0002] We speak of very low magnetic field MRI when the stationary magnetic field polarization of nuclear spins has an intensity between approximately 50 pT and 100 mT, corresponding to Larmor frequencies on 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 document WO 2022 / 117969 that a system based on SQUIDs (Superconducting Quantum Interference Devices), and more particularly low-temperature direct current SQUIDs (dc-SQUIDs), is used for the detection of such a weak MRI signal. Further information on SQUIDs and their use is provided by (Fagaly) and (Clarke, Braginski).

[0004] As illustrated in [Fig. 1], a DCS DC SQUID is in the form of a superconducting loop BS interrupted by two Josephson junctions JH, JJ2. When a direct current I flows through the loop, the potential difference V across the Josephson junctions is a sinusoidal function of the magnetic field flux <t>s through the loop. Typically, a bias coil (BP) is used to fix 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 ip flowing through the bias coil so as to keep the voltage Vs supplied to the FLL constant after being amplified by a low-noise amplifier (LNA). In this case, the current ip—or another signal generated by the FLL and proportional to ip—can be taken as the output signal providing a measure of the flux <FS. En variante, si le SQUID présente une dynamique suffisante, la boucle FLL peut avoir une bande passante inférieure à la fréquence du signal utile ; dans ce cas, le signal de sortie est constitué par une variation de tension haute fréquence aux bornes du SQUID, VSHF.

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

[0006] It is known that SQUIDs can be fabricated with both "conventional" low-critical-temperature superconductors and high-critical-temperature superconductors. A "low critical temperature" is defined as a critical temperature of 20 K or less. In the aforementioned WO 2022 / 117969, the use of low-critical-temperature SQUIDs—maintained at a temperature well below the transition temperature, for example, on the order of 4 K—is recommended due to their lower noise level, greater ease of fabrication, and better figure of merit RnIc (where Rn is the junction resistance in the normal state and Ic is its critical current) compared to high-critical-temperature SQUIDs. Furthermore, it is known that low-temperature superconducting materials can exhibit superior mechanical, shaping, and stability properties compared to high-temperature superconductors.

[0007] The superconducting loop BS has a small surface area (a few pm² or tens of pm²), and therefore intercepts a weak magnetic flux. Furthermore, since it must be kept at cryogenic temperatures, it cannot always be optimally positioned to detect a magnetic field, for example, near the body of a patient undergoing an MRI scan. For these reasons, it is known to use what is called a flux concentrator CF. The flux concentrator CF comprises a first conductive coil BC, called the capture coil, connected in series with a second conductive coil BE, called the input coil. The latter is generally made using planar technology and is 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 denoted by MES.The BC pickup coil, which effectively acts as an antenna, has a significantly larger effective area—the product of the number of turns and the geometric area of ​​each turn—than that of the superconducting loop. Consequently, it carries a greater flux, which is transferred to the SQUID's superconducting loop via the input coil. Furthermore, the pickup loop does not need to be maintained at cryogenic temperatures and can therefore be positioned more freely.

[0008] The BA pickup coil 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' and BG1” are planar, located in planes perpendicular to a z-axis and separated by a distance d (called the “baseline”) along said axis. It is easy to understand that the current induced in the gradiometric coil BG1 depends on the variation of the Bz component of the magnetic field along the baseline d, which—to a first approximation—is proportional to the intensity of a The magnetic field decreases as the cube of its distance from the source, but the magnetic field gradient decreases as the fourth power of this distance. Therefore, replacing a simple coil with a gradiometric coil helps to attenuate the impact of distant noise sources. Figure 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, to a first approximation, on the second derivative of Bz with respect to z: Similarly In this way, it is possible to make higher order gradiometric coils, with the rejection of noise from distant sources increasing with said order.

[0009] Figure 2C shows a volumetric "horsesaddle" coil (BSC), used in MRI because of its field homogeneity properties. Figure 2D shows the gradiometric (first-order) version of a horsesaddle coil (BSCG), consisting of two horsesaddle sub-coils: a first, internal coil, having—in the example shown—two turns of wire, and a second, external, and larger coil, having—again, in the example shown—a single turn of wire. The difference in the number of wire turns allows the two sub-coils to have the same inductance despite their different sizes. The windings are wound in such a way that the current flows in opposite directions in the two sub-coils.

[0010] One drawback of these gradiometric setups is that the antennas use long lengths of wire due to the succession of coils compensating each other. The receiving antennas thus constructed generally have a relatively high inductance, which can pose a problem for detection sensitivity with SQUIDs. Indeed, for quantum interference to be observed, the self-inductance Ls of a SQUID must satisfy the condition

[0011] H) 2ï; > ^k B T

[0012] where ¢0 is the flux quantum, T the absolute temperature and kB the Boltzmann constant. Typically, Ls is therefore on the order of 1 pH.

[0013] If we call <hBC le flux à travers la bobine de captation BC et LBC son inductance, le courant iBC qui parcourt ladite bobine et la bobine d’entrée BE vaut

[0014]

[0015]

[0016]

[0017] • _ (2) ^BC — Lbc+Le The Osa flux, through the superconducting loop of the SQUID, is ^sa ~ MESiBC (3) AveC MES = k^LELs (4)

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

[0019] k^Ls (5) &SA ~ Lbc+Le ^BC

[0020] It follows from equation (5) that, for a given OBC flux through the capture coil, the Osa flux seen by the SQUID is maximum for LE = LBC. However, to maximize OBC, it would be desirable to take LBC to be large, while LE must be of the same order of magnitude as Ls 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 to it inductively efficiently). The low value of the SQUID's inductance therefore imposes a maximum limit on that of the input coil, and thus on that of the capture coil. It follows that LBC cannot exceed a few pH or tens of pH, which is very restrictive, particularly for the construction of volume coils in an MRI scanner, which, due to their dimensions, tend to exhibit high inductance values.More specifically, it is preferable that the LE / LBC ratio be within a range between 0.4 and 2.5 so that the sensitivity is at least equal to 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 portion of which is dedicated to capturing the signal of interest, the remainder being dedicated to noise suppression. Consequently, a large part of the inductance of the capturing coil—the value of which is limited by the matching constraints described in the preceding paragraph—is not used to collect the signal of interest. This problem is even more serious for higher-order gradiometers.

[0022] In the aforementioned WO 2022 / 117969 document, a transformer is interposed between the pickup coil and the input coil. This transformer increases the impedance LA while respecting the matching requirements. However, it exhibits significant losses, especially when the transformation factor is high, which limits the usefulness of this solution. Furthermore, transformers operating at low temperatures can hardly use ferromagnetic cores.

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

[0024] According to one aspect of the invention, this objective 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 the 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 enabling the acquisition of a gradiometer of order N+1.

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

[0026] An object of the invention is therefore a radio frequency gradiometer comprising two conductive coils, called capture 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 make 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 radio frequency magnetic field values ​​corresponding to the two capture coils.

[0027] According to particular embodiments of such a radio-frequency gradiometer:

[0028] - Said sensing coils can be configured to operate in a frequency range between 1 kHz and 100 MHz.

[0029] - Said SQUID magnetometers may include current-carrying SQUIDs continuous.

[0030] - Said SQUID magnetometers may include low-frequency SQUIDs critical temperature.

[0031] - Said electronic processing circuit may be an analog circuit comprising an adder or subtractor amplifier to add or subtract said output signals. More specifically, each said SQUID magnetometer may include a low-noise amplifier to generate the corresponding output signal by amplifying a voltage drop across the terminals of a respective SQUID.

[0032] - Each said magnetometer may include a flux-locking loop.

[0033] - Each said pickup coil may have an inductance greater than or equal to at pH 10.

[0034] Another object of the invention is a magnetic resonance imaging device comprising at least one radiofrequency gradiometer according to any one of the preceding claims, used as a receiving antenna. According to particular embodiments:

[0035] - Said capture coils can form volumetric antennas.

[0036] - The device may also include a coil for generating a magnetic field stationary polarization of nuclear spins, with an amplitude between 10 pT and 100 mT.

[0037] Other features, details and advantages of the invention will become apparent from the description given with reference to the accompanying drawings provided by way of example, which represent, respectively:

[0038] [Fig.1], described above, a SQUID magnetic field detection system known from the prior art.

[0039] [Fig.2A], [Fig.2B], [Fig.2C] and [Fig.2D], described above, different types antennas that can be used with the detection system of [Fig.1] or for the implementation of the invention;

[0040] [Fig.3], a radio frequency gradiometer according to a first embodiment of the invention;

[0041] [Fig.4], a radiofrequency gradiometer according to a second embodiment of the invention; and

[0042] [Fig.5], a nuclear magnetic resonance imaging (scanner) device using at least one radiofrequency gradiometer according to an embodiment of the invention.

[0043] The apparatus of [Fig. 3] comprises two SQUID magnetometers analogous to that of [Fig. 1], MSA, MSB. Each of these magnetometers comprises a DC SQUID DCSA, DCSB inductively coupled to an input coil BEA, BEb and to a biasing coil BPA, BPb which sets its operating point and is driven by a flux-locked loop (not shown). The input coil is connected in series with a pickup coil BCaBCb. These two pickup loops are arranged like the two windings of the gradiometric antenna of [Fig. 2A]: they are planar, arranged in planes perpendicular to the z-direction (the "axis of sensitivity") and therefore sensitive to the Bz component of an external magnetic field; they are also separated by a distance d in said z-direction. If z0 is the coordinate along the z-axis of the midpoint of the two coils, coil BCA is at position z0+d / 2 and coil BCb is at position z0+d / 2.

[0044]

[0045]

[0046]

[0047]

[0048]

[0049]

[0050]

[0051]

[0052]

[0053]

[0054] The voltages across the two SQUIDs, VA and VB, are amplified by low-noise amplifiers, LNAa and LNAB. The winding directions of the pickup coils, BCa and BCb, and the input coils, BEA and BEb, are chosen such that a magnetic flux through the pickup coils induces VA and VB voltages of opposite signs. For example, the pickup coils may have windings wound in opposite directions and the input coils in the same direction, or vice versa. The VSA and VSB outputs of the low-noise amplifiers are fed into 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 Rb, R2, R3, and R4. If Ri = R2j, the output voltage Vs of the operational amplifier OA is proportional to VA + VB, and therefore to the difference in magnetic fluxes across the two pickup coils. The system thus 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 for the acquisition of a useful signal; this architecture therefore allows for a reduction in the constraints on the value of this inductance. In more detail, if LBC=LE (we consider that the two pickup coils BCa, BCb have the same inductance LBC, the two input coils BEA, BEb the same inductance LE and the two SQUIDs DCSA, DCSB the same inductance Ls) equation (5) becomes Or <t>BC is equal to (pA= cp(zQ+j ) or 0£ = ^^o_^ depending on whether we consider one or the other of the two capture reels. The voltage VA across the terminals of DCSA is therefore = 0(z„ + 4 ) <8) Where Ks is the flux-voltage transfer coefficient of the SQUID (assumed to be the same for DCSa and DCSb). Similarly, the voltage VB across DCSB is

[0055] By introducing a parameter A which takes into account the amplifiers LNAa, LNAB (assumed to be identical), the effect of the flux-locked loop and the gain of the voltage adder circuit - equal to 1+R3 / R4 - we obtain:

[0056] [¢(¾ A )-♦M ) ] dg |. =SK.AÆ,^ d % <10>

[0057] In equation (10), the last equality was obtained by considering that the two pickup coils have the same effective area SBc (in the simple case of a planar coil, this is the area projected onto the xy plane perpendicular to the z-axis and multiplied by the number of windings) and that therefore <e>=BzSbc.

[0058] In the embodiment of [Fig. 4], the winding directions of the pickup coils BCa, BCb and the input coils BEA, BEb are chosen such that a magnetic flux through the pickup coils induces voltages VA, VB 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 Vs is proportional to

[0059]

[0060] As in the case of [Fig.3], the system therefore constitutes a first-order gradiometer, identified by the reference GR'.

[0061] Figure 5 schematically represents a magnetic resonance imaging (MRI) device (or "scanner"). Such a device includes a BCM coil for generating a uniform static magnetic field B0 oriented along a longitudinal direction, which allows the nuclear spins of the atoms in a sample—for example, a part of a human body, or the entire body—to be oriented (polarized) in that direction. Gradient coils (not shown) apply non-uniform static magnetic fields, necessary for obtaining spatial resolution. Radio-frequency coils (not shown) apply excitation pulses to the nuclear spins, after which the nuclear spins de-excite by emitting radio-frequency radiation.One or more gradiometers (GR) 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.

[0062] The invention has been described with reference to certain embodiments, but several variations are possible. In particular:

[0063] - Other types of pickup coils can be used, such as coils volumetric. For example, the pickup coils can be the sub-antennas. of a gradiometric saddle-type volume antenna, of the type shown 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.

[0064] - Although it is generally preferable for both antennas to have the same Inductance is also a factor; it is possible that the two elementary magnetometers may have different inductances. Similarly, the SQUIDs and the input coils of the two magnetometers may not be identical. The resulting differences in output signal intensity must then be compensated for by the electronic processing circuit. In the embodiment shown in [Fig. 3], this is achieved by choosing different values ​​for the resistors RI and R2. Likewise, it is possible to compensate for any difference in effective area.

[0065] - The electronic processing circuit can add or subtract the signals that it receives input (where appropriately weighted to take into account a possible difference between the two elementary magnetometers), depending on the winding directions of the capture coils and the input coils.

[0066] - The electronic processing circuit may have a different topology from those described. For example, the processing can be partially or totally digital.

[0067] - Although the use of SQUID at low critical temperatures is preferred, that of High critical temperature SQUID is not excluded.

[0068] - In some cases, the flux-locking loop can be omitted, the polarization with the SQUIDs being fixed. Conversely, the output signal of each elementary magnetometer can be provided by the flux-locking loop instead of being taken from the terminals of the SQUIDs.

[0069] - The invention can be implemented in other frequency ranges and for Other applications besides MRI – for example, magnetic resonance spectroscopy, mineral exploration, etc. In all cases, the magnetic field polarizing nuclear spins may be outside the intensity range indicated as preferred. References

[0070] (Clarke, Braginski) John Clarke, Alex I. Braginski (editors) The SQUID Handbook, Vol. I (2004)

[0071] (Fagaly) RL Fagaly “Superconducting quantum interference device instruments and applications” Review of Scientific Instruments 77, 101101 (2006)

[0072] (Matlashov) A. Matlashov et al. “Electronic Gradiometer using HTC SQUIDs with Fast Feedbeck Electronics” IEEE Transactions on Applied Superconductivity, Vol. 11, No. 1, March 2001.< / e> < / t> < / t>

Claims

Demands

1. 1. Magnetic resonance imaging (MRI) apparatus comprising at least one radio frequency gradiometer (GR, GR') comprising two conductive coils (BCA, BCb), referred to as 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 (BEA, BEb), referred to as input coil, to make a flux concentrator (CF), each of said input coils being inductively coupled to a SQUID magnetometer (MSa, MSb); characterized in that the gradiometer also comprises an electronic processing circuit (EPC) configured to linearly combine output signals (VSA, VSB) from said SQUID magnetometers to obtain a signal (Vs) representative of a difference between radio frequency magnetic field values ​​(Bz) corresponding to the two capture coils.

2. 2. Magnetic resonance imaging apparatus_according to claim 1 in which said capture coils (BCA, BCb) are configured to operate in a frequency range between 1 kHz and 100 MHz.

3. 3. Magnetic resonance imaging apparatus according to any one of the preceding claims wherein said SQUID magnetometers (BCa, BCb) comprise DC current SQUIDs (DCSa, DCSb).

4. 4. Magnetic resonance imaging apparatus according to any one of the preceding claims wherein said SQUID magnetometers (BCa, BCb) comprise low critical temperature SQUIDs (DCSA, DCSB).

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

6. 6. Magnetic resonance imaging apparatus_according to claim 5 in which each said SQUID magnetometer comprises a low noise amplifier (LNAa, LNAB) for generating the corresponding output signal by amplifying a voltage drop (VA, VB) across the terminals of a respective SQUID (DCSA, DCSB).

7. 7. Magnetic resonance imaging apparatus_according to any one of the preceding claims in which each said magnetometer (MSa, MSb) comprises a flux-locked loop (FLL).

8. 8. Magnetic resonance imaging apparatus according to any one of the preceding claims wherein each said capture coil (BCA, BCb) has an inductance greater than or equal to 10 pH.

9. 9. A magnetic resonance imaging (MRI) apparatus according to any one of the preceding claims, wherein said pickup coils (BCA, BCb) form volume antennas.

10. A magnetic resonance imaging (MRI) apparatus according to any one of the preceding claims, comprising a coil (BCM) for generating a stationary magnetic field (B0) for polarizing nuclear spins, with an amplitude between 10 pT and 100 mT.

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