JPH0566811B2 - - Google Patents

Description

DETAILED DESCRIPTION OF THE INVENTION The present invention relates to a method for generating a motion signal in an MRI tomography apparatus depending on the motion of a subject.

The invention also relates to an MRI tomography apparatus for carrying out such a method, comprising a first radio-frequency coil system for generating a radio-frequency magnetic field having the Larmor frequency and for receiving spin resonance signals.

MRI tomography devices of this type are known. And generally, the same high-frequency coil system is used both for generating the high-frequency magnetic field and for receiving the spin resonance signal. Although the term "high-frequency coil system" is used here, this term should be interpreted broadly. For example, resonators such as those described in German Patent Application No. P3347597 must also be included in this understanding. In this way, a high frequency magnetic field is generated, the frequency of which falls within the so-called Larmor frequency range. Larmor frequency is
It is proportional to the strength of the static magnetic field generated by an MRI tomography device, and reaches approximately 42.5MHz/T in the case of hydrogen. The high frequency magnetic field excites each spin within the subject. Therefore, after this magnetic field disappears, a spin resonance signal is induced in the high frequency coil system as a quasi-echo. This signal is further processed by a computer to determine the nuclear spin density and relaxation time T1 within the specimen.
T2 or one of the two is required.

When an object is examined many times using such an MRI tomography device, a high-frequency magnetic field is generated during the nuclear cycle, and resonance signals are received.
After a resonant signal is received, a relatively long time (several
(on the order of magnitude of 100ms). The total inspection time is therefore also relatively long. Generally it takes more than 1 minute. Obviously, it is unavoidable that the patient moves during such a long period of time, especially breathing and swallowing movements. However, these movements tend to falsify test results. To avoid this phenomenon, use the known
MRI tomography machines are also equipped with motion detectors that detect patient movement, but these motion detectors can be thermally, pneumatically, mechanically, electrically, or optically operated and must be attached to the patient. It won't happen. The motion signals thus generated must then be transmitted via separate channels. Nevertheless, the motion signals thus generated can significantly reduce the moving artifacts caused by patient motion.

One way to take this into account is to start each cycle only when the movement signal indicates that the patient's body will again assume a defined position (so-called triggering). Another method is to complete the cycle regardless of the motion signal from the motion detector, preferably at regular time intervals, which eliminates spin resonance signals from the reconstructed image that occur when the object is in motion (so-called gating). The discarded cycle must now be repeated in the same manner.
It is also possible to combine both methods, ie gating and triggering.

An object of the present invention is to provide an MRI tomography device that can easily detect motion.

To achieve this, in an MRI tomography apparatus for generating motion signals, a first high-frequency coil means surrounding the subject also generates a high-frequency magnetic field at the Larmor frequency and receives a spin resonance signal from the subject. , a second high-frequency coil means disposed adjacent to the subject and inside the first high-frequency coil means and for generating a motion signal.
It is characterized in that it comprises a high frequency coil means and an impedance measuring means for measuring the impedance of the second high frequency coil means for determining the motion signal.

In the present invention, the Q and stray capacitance of the high frequency coil system are
It is used to change in response to all movements of the patient placed within the magnetic field of this radio frequency coil system, such as breathing, swallowing, heartbeat and peristalsis. thus,
The impedance of this coil system varies in magnitude and phase. According to the invention, this impedance is measured at least at intervals during the test, and a motion signal is derived from the resulting measurement signal, for example by rectification. Each value of impedance is associated with a given phase of motion.

Two methods exist to implement this method. The first method is to generate a high frequency magnetic field,
It is based on an MRI tomography device that is equipped with an impedance measurement unit that measures the impedance of a high-frequency coil system used to receive spin resonance signals. The method is characterized in that the impedance measuring unit is activated during the examination of the subject and the measuring signal thus generated is used as the motion signal.

It should be noted that it is necessary to measure the impedance of the high-frequency coil system of the MRI tomography device. This is because placing a patient in the magnetic field of the high-frequency coil system affects the Q or conformity and resonance frequency of the high-frequency coil system. Therefore, an impedance measuring unit is provided which acts on the adjustment member for automatically adapting or readjusting the high-frequency coil system. After starting the actual test, the impedance measuring unit is no longer activated. This can also be used to generate motion signals.

In a second method, the MRI tomography device comprises a second high-frequency coil system for generating motion signals;
The impedance of this second high frequency coil system is measured by an impedance measuring unit.

The first method is simpler. This is because only one high-frequency coil system is required and it performs dual functions. That is, it serves on the one hand to generate high-frequency magnetic fields and to receive spin resonance signals, and on the other hand to generate motion signals. However, a second method provides a second coil system and is structured to respond much better to the motion of the object. In addition, this allows continuous measurement during the entire inspection. The examination is not adversely affected by the formation of motion signals, especially if the frequency is shifted considerably from the Larmor frequency.

The invention will be explained in detail with reference to the drawings.

As shown in FIG. 1, the MRI tomography device comprises an electromagnet consisting of four coils 1 generating a strong, uniform static magnetic field extending in the direction of a common horizontal coil axis. A patient 3 sitting on a bed 2 and inside an electromagnet is surrounded by a high-frequency coil 4. This high frequency coil 4 generates a pulsed high frequency magnetic field that extends perpendicularly to the main magnetic field generated by the electromagnet. The frequency of the high-frequency magnetic field is proportional to the magnetic flux density of the main magnetic field,
It can be 0.1T to 4T depending on the structure of the electromagnet.
This proportionality constant corresponds to the gyromagnetic ratio (approximately 42.5M
Hz/T). In this way, nuclear spins are resonantly excited by the high-frequency magnetic field in the space surrounded by the high-frequency coil.

The MRI tomography device also comprises four gradient coils 5. These gradient coils 5 generate a magnetic field that extends in the direction of the main magnetic field and varies linearly in this direction. The MRI tomography device also comprises further gradient coils which generate magnetic fields varying in two directions, again extending in the direction of the main magnetic field, but perpendicular to this direction. When these gradient coils are energized, a high frequency magnetic field is generated followed by the high frequency coil system 4.
The phase of the signal induced within is varied depending on the nuclear spin density within the object region being examined.
Therefore, in principle, the nuclear spin density within a two-dimensional or three-dimensional region of an object can be determined by such an MRI tomography apparatus.

FIG. 2 shows a block diagram of a first embodiment according to the invention. This high frequency coil system 4 is accompanied by a variable capacitor 7 to form a parallel resonant circuit. This parallel resonant circuit is connected to ground on the one hand and to a switch 9 via another variable capacitor 8 on the other hand.
When the switch 9 is in the position shown, the circuitry 4, 7, 8 receives power from a high power radio frequency transmitter 10 whose carrier frequency corresponds to the spin resonance frequency. When the switch 9 is in the position not shown, the preamplifier 11 of the receiver (not shown) is connected to the network formed by the variable capacitors 7 and 8 and the high-frequency coil 4, and the preamplifier 11
receives the signal induced in the high frequency coil 4 by spin resonance.

At the spin resonance frequency, the input impedance of the preamplifier 11, the output impedance of the high-frequency transmitter 10 and the impedance of the network 4, 7, 8 are equal, for example 50Ω. After introducing the patient 3 and before actually starting the test, capacitors 7 and 8 are adjusted, i.e. preferably automatically using an impedance measuring device (not shown) to adjust this matching condition. . However, as the patient moves within the coil, ie, breathing, swallowing, heartbeat, and peristalsis, this alignment changes depending on the position of the coil. This is because the Q and stray capacitance of the coil 4 are affected. Therefore, network 4,
The instantaneous value of the impedance 7,8 is a measure of the current phase of the motion of the object being examined, and a motion signal can therefore be derived from this.

For this purpose, a reflectometer 12 is connected between the transmitter 10 and the switch 9 and the output signal of the reflectometer 12 is processed in a reflectometer receiver 13 . Receiver 1
A motion signal appears at the output terminal 14 of 3. As long as the radio frequency transmitter 10 and the circuitry 4, 7, 8 are matched, the output signal of the reflectometer 12 will be approximately zero. However, if they are not matched, a portion of the applied high frequency power will be reflected. The output of the reflectometer then carries a high frequency signal whose amplitude depends on the reflection factor, ie the degree of mismatch. This high frequency signal is reflected and amplified by the receiver 13 and then appears at the output terminal 14. This signal is then compared with two presettable thresholds corresponding to a predetermined impedance and therefore a predetermined movement. A transition between these two thresholds can be used to initialize the control means.

FIG. 5 is a typical time diagram of an examination performed by such an MRI tomography apparatus. A so-called 90° pulse is generated during the period Ta.
That is, a high-frequency transmitter 10 is connected to the high-frequency coil 4 via a switch 9, so that the nuclear magnetization of the object being examined is tilted at exactly 90° to the direction of the main magnetic field. Next, usually one or more so-called 180° pulses Tb are applied, after which the signal induced in the coil is received (period Tc). At this time, the switch 9 assumes a position different from that shown. The typical duration of this procedure is approximately 100ms. This procedure is repeated regularly with a period of 600 ms, which is long compared to the duration of the procedure itself. The magnetic field of the gradient coil is then switched in a defined manner during each repetition.

In the device shown in FIG. 2, impedance measurement is possible only during the period when the coil 4 receives power from the high frequency transmitter 10. Therefore, in FIG. 5 these are periods Ta and Tb. Assuming that the patient's body is in approximately the same phase during relatively closely consecutive periods Ta...Tc, so-called gating can be performed using the derived motion signal. When the motion signal is within a predetermined amplitude range, ie when the patient's body is in a substantially constant position, the spin resonance signal attracted to the coil 4 during the period Tc is evaluated. If this is not the case, the induced signal is not evaluated and the measurement has to be repeated (with the same gradient field).

In the embodiment shown in Figure 2, if triggering processing must also be performed, the last pulse of one measurement is
The impedance must be measured during a relatively long period that elapses between the reception of Tc and the delivery of the first pulse Ta of the next measurement. Therefore, the high frequency transmitter 10 must also apply a signal to the coil 4 during this period. However, in order to avoid that these measurements influence the spin relaxation that occurs in the period between Tc and Ta, the lever signal must be significantly reduced in power or performed within well-defined time intervals within this period. You have to add it like this. In the latter case, if measurement is not performed at a frequency other than the Larmor frequency, the measurement is performed in a considerably shorter period compared to the periods Ta and Tc. However, in this case it is more difficult to generate a motion signal. However, the reflected signal must have a relatively small amplitude or be formed within a relatively short time interval.

As previously discussed, patient movement affects the coil's Q and inductance. For this reason, the circuit network 7,
8,9 resonance curves shift towards higher or lower frequencies. However, there is no clear relationship between impedance and the associated phase of motion.
This can be avoided by slightly detuning the networks 4, 7, 8 with respect to the measurement frequency (usually the spin resonance frequency). This detuning must be small relative to the 3dB bandwidth of the network. However, it must be large so that the resonant frequency of the network 4, 7, 8 always remains below or above the measurement frequency during any phase of motion. A suitable detuning value is a 3dB bandwidth of 300KHz and a spin resonance frequency of approximately 85MHz.
If it is 20KHz.

FIG. 3 shows another embodiment of the invention. Corresponding parts are given corresponding symbols. switch 9
is directly connected to the high frequency transmitter 10. In contrast to the switch shown in FIG. 2, this switch has three positions. In one position, the power of the radio frequency transmitter 10 is applied to the network 4, 7, 8. In the second switch position, an impedance measuring unit is connected to this network. The impedance measuring unit consists of another high frequency generator 15 and an impedance measuring bridge 16 which is supplied with power from this high frequency generator 15.

The impedance measurement bridge 16 is, for example,
Configured as a Wheatstone bridge, the impedance of the network 4, 7, 8 is connected to the high frequency generator 15.
It is designed to be balanced against a predetermined reference impedance, for example 50Ω, required for matching at the frequency of . The power of the high frequency generator 15 is considerably lower than the power of the high frequency transmitter 10. The output frequency of the high frequency generator 15 is shifted from the output frequency of the high frequency transmitter 10.
The 3 dB bandwidth of 8 is proportionally small and large enough that nuclear spins in the region of the object are not thereby excited. In this way, the actual test is not influenced by the impedance measurement. It is also advantageous to offset the resonant frequencies of the networks 4, 7, 8 from the measuring frequency of the high-frequency generator 15.

By using such a circuit, impedance measurement can be performed at any time between periods Ta and Tc, between Tb and Td, and between Tc and Ta. That is, impedance measurements can be carried out whenever the high frequency coil 4 does not generate a magnetic field or a resonant signal is induced therein. The motion signal generated by the impedance measuring unit can thus be used not only for gating but also for triggering.

The signal appearing at the output of the impedance measuring bridge usually requires further processing (amplification, rectification). Appropriate processing equipment is required for this purpose. However, as shown by the dashed line, the output signal of the impedance measuring bridge 16 is instead transferred to the preamplifier 1.
It is also possible to amplify one or a portion thereof. In that case, the preamplifier, impedance measuring unit 1
5, 16 (balancing the circuit networks 4, 7, 8 before testing and generating motion signals during testing)
and a radio frequency coil (detection of motion and excitation and reception of spin resonance signals) performs a dual function.

FIG. 4 shows an embodiment of the impedance measuring units 15, 16 shown in FIG. 3, which is easy to make. A high frequency generator 15 (one of which is grounded) powers an inductive RF bridge circuit comprising four inductances 17, 18, 18a and 19. Inductances 17 and 19 are equal, and so are inductances 18 and 18a. The two inductances 17 and 19 are also permanently magnetically coupled. A transmitter is thus obtained. A resistor 20 is connected to the ground points of the inductances 17 and 18. The other terminal of resistor 20 is grounded. The resistance value of resistor 20 is RF coil system 4
Correspond to the standard impedance (50Ω). When measuring the impedance at the connection point of the inductances 19 and 18a when the impedance is exactly the value (50Ω) required to match the frequency of the high-frequency generator 15, the circuit network 4, 7, Connect to 8. At this time, voltages of equal magnitude and opposite phase appear at both connection points of the inductance. Therefore, no voltage is generated at the center tap 21 of the inductance. In the unmatched condition, the voltages appearing at both junctions of the inductances are not equal and therefore a voltage different from essentially zero appears at the center tap 21. This voltage is a measure of mismatch.

So far, we have assumed that the coil that generates the high-frequency magnetic field to excite the nuclear spins also receives the spin resonance signal. However, the present invention can also be applied when exciting a magnetic field and receiving a spin resonance signal are performed using separate coils. At this time, the embodiment shown in Fig. 2 is used in combination with a coil necessary to generate a high-frequency magnetic field, and the embodiment shown in Fig. 3 is used in combination with a coil whose purpose is to receive a spin resonance signal. be able to.

It has also been assumed up to now that the impedance measurement is performed at a frequency that is at least close to the frequency of the spin resonance signal. However, this measurement can also be carried out at a frequency of the second order of the resonance frequency of the networks 4, 7, 8. However, the impedance of this network varies significantly in magnitude and phase even at the secondary resonant frequency. This gives the advantage that impedance measurements do not affect the excitation of nuclear spins.

In the embodiments described so far, it has been assumed that the radio-frequency coil system serves on the one hand to excite the nuclear spins and on the other hand to form a motion signal. In the following, an embodiment will be described in which a separate high-frequency coil system is provided for generating the motion signal.

FIG. 6 is a cross-sectional view of a portion of an MRI tomography apparatus including such a separate radio frequency coil system. Inside the circular area 30 corresponding to the free opening of the electromagnet 1 (Fig.
A high frequency coil system 4 suitable for high frequencies is provided as described in No. This two-part coil system supplies current in the conductors of the upper loop (which extends perpendicular to the plane of the drawing) in the opposite direction to the corresponding conductors of the lower loop. In this way, the high frequency magnetic field generated by this coil extends perpendicularly to the static magnetic field in the X direction.

The high-frequency coil system 4 is attached in a manner not shown to a hollow cylindrical plastic body 31 which is rigidly connected to the bed on which the patient rests. A high-frequency coil system 33 serving to generate the motion signal is connected to the plastic body 31 via a support 32. This coil can also be displaced in the y or z direction. It can also be connected to the plastic body in a manner that is invisible to the patient. For example, on the outside of the plastic body. The coil has one or more turns and is arranged so that the magnetic field generated by the coil extends in the y direction. High frequency coil system 4
If the plastic body 31 with the Arrange it so that it extends in the direction.

It is only important that the magnetic field generated by the high-frequency coil 33 extends essentially perpendicular to the magnetic field generated by the high-frequency coil system 4. This is because the high-frequency coil system 4 and the high-frequency coil 33 are not highly magnetically coupled to each other, and therefore, on the one hand, the magnetic field generated by the high-frequency coil system 4 becomes smaller due to the presence of the high-frequency coil 33. On the other hand, the signal induced in the high-frequency coil 33 by the high-frequency coil system 4 also becomes smaller.

Reference numeral 34 schematically represents another, preferably flat, coil. This coil is connected to the patient's body 3
, and its magnetic field is also essentially oriented in the y direction. This coil is made in the same way as a surface coil used in MRI tomography equipment solely for the reception of spin resonance signals (rather than for the excitation of nuclear spins). In contrast to coils 33 which are fixed to the plastic body, the structure of such surface coils can be adapted to the relevant application. This fact, and the fact that the coil can move in close proximity to the area whose motion is to be monitored, provides an inherently high sensitivity. Such a coil is therefore able to extract useful motion signals even from relatively small movements, such as, for example, respiratory movements. On the other hand, these coils need to be attached to the patient's body before the examination.

FIG. 7 shows a circuit for deriving the motion signal. The circuit comprises an impedance measuring unit, for example an impedance measuring bridge 35, in one branch of which a high frequency generator 4 is provided.
6 is connected, and high frequency coils 33 or 34 are connected to the other branches via a conversion circuit network. The frequency of the radio frequency transmitter is significantly higher than the Larmor frequency, which is derived from the strength of the magnetic field of the magnet in the case of hydrogen. If the magnetic field strength of the magnet is, for example, 2T, the Larmor frequency is approximately 85MHz, and the frequency of the high frequency transmitter must be between 100MHz and 200MHz, and must not match the high frequency of the Larmor frequency. It won't happen. If this MRI tomography device is also used to investigate the distribution of elements other than hydrogen, such as sodium, phosphorus, or fluorine, the Larmor frequencies for sodium, phosphorus, and fluorine at 2T (approximately 68 MHz, 35 MHz, and 80MHz)
It must not coincide with the high frequency of

The desired impedance of the high frequency coil is 50.
The conversion into a real resistance of Ω is achieved by, on the one hand, a tuning capacitor 36 connected in series with a parallel resonant circuit 37 in parallel with the high-frequency coils 33, 34;
and on the other hand by the trimmer capacitor 38. One end of the coils 33, 34 is attached to this trimmer 3.
8 and a parallel resonant circuit 39 connected in series with it.
and to one branch of the impedance measuring bridge 35. Parallel resonant circuits 37 and 39 are tuned to the Larmor frequency and the high frequency transmitter 3
6 are designed such that their capacitive impedance at the frequency is lower (at least not substantially higher) than the capacitance of capacitors 36 and 38. Therefore, the capacitive impedance of the series or parallel branches can be varied by varying the capacitance of capacitors 36 and 38.

After putting the patient, and maybe the high frequency coil 34
After fixing and before starting the test, adjust capacitors 36 and 38 (preferably automatically) so that the input resistance of the network formed by high-frequency coils 33, 34 and elements 36...39 is equal to that of the high-frequency transmitter. 46
A predetermined value, for example, 50Ω, is set at the frequency of . The input resistance is thus matched to the impedance measuring bridge. During the next inspection,
Output terminal 21 of impedance measurement bridge 35
is responsible for the measurement signal whose amplitude has the value zero or a minimum value in the matching state. In response to patient movements, such as breathing, swallowing, heartbeat and peristalsis, the Q and stray capacitance of the coil 33 or 34 changes, thus changing the impedance of the circuitry connected to the impedance measuring bridge 35; The amplitude of the output signal increases with the phase of motion involved. The signal at the output terminal 21 can be used as a motion signal after being preferably rectified by a circuit with a small time constant.

Generally, a high frequency coil 33 is used for generating a motion signal by a magnetic field generated by the high frequency coil system 4.
and 34 cannot be sufficiently prevented from inducing voltages. This is especially true when the position relative to the high-frequency coil system 4 can change, as is the case with the high-frequency coil 34. As a result of kneading, a current flows through the circuit connected to the high frequency coil 33 or 34, a load is applied to the high frequency coil system 4, and the magnetic field generated by the high frequency coil system 4 is distorted. Also, the signal appearing at output terminal 21 is prone to errors. This fact must be taken into account when the power applied to the high-frequency coil system 4 is considerably greater than the power applied to the high-frequency coil system 33 or 34. These two effects are reduced by the resonant circuits 37 and 39, which have high impedance at the frequency of the high frequency coil system 4.

Elements 36 and 37 are high frequency coils 33 or 34
operate in parallel resonance at the frequency of the high frequency generator 46. However, when the high-frequency coils 33 and 34 operate in series resonance due to the trimmer 38 being connected in series, the series branch connected in parallel with the parallel resonance circuit 37 can be omitted.
This is because the series resonant circuit already formed at this time has high impedance at frequencies far from the series resonant frequency, such as the frequency of the high-frequency coil system 4.

An advantage of the embodiment with a separate high-frequency coil system for generating the movement signal is that the state of movement of the patient is determined continuously during the entire or almost the entire examination. In addition, the high frequency coil 33 or 34
The magnetic field generated by this does not excite nuclear spins. This is because the frequency of the high-frequency transmitter deviates considerably from the Larmor frequency.

[Brief explanation of the drawing]

Figure 1 is a schematic longitudinal cross-sectional view of an MRI tomography device, and Figures 2 and 3 show a type of MRI tomography in which a high-frequency coil system that excites nuclear spins and also receives resonance signals is also used to generate motion signals. 4 is a circuit diagram of the impedance measuring circuit used in the embodiment shown in FIG. 3, FIG. 5 is a time diagram showing the operation of the high frequency coil system, and FIG. 6 is a block diagram of two embodiments of the device. The figure includes a separate high-frequency coil system for forming the motion signal.
FIG. 7, which is a cross-sectional view of the MRI tomography apparatus, is a circuit diagram of an impedance measurement circuit for such an MRI tomography apparatus. 1... Coil, 2... Bed, 3... Patient, 4... High frequency coil, 5... Gradient coil, 7... Variable capacitor,
8... Variable capacitor, 9... Switch, 10... Transmitter, 11... Preamplifier, 12... Reflectance meter, 13...
Reflection measurement receiver, 14... Output terminal, 15... High frequency generator, 16... Impedance measurement bridge,
17-19...Inductance, 20...Resistance, 21
...Central tap, 30...Circular area, 31...Hollow cylindrical plastic body, 32...Support, 33...High frequency coil (for motion signal), 34...Another coil, 35
... Impedance measurement bridge, 36... Tuning capacitor, 37... Parallel resonant circuit, 38... Trimmer, 3
9...Parallel resonant circuit, 46...High frequency generator.

Claims (1)

[Scope of Claims] 1. In an MRI tomography apparatus for generating motion signals, first high-frequency coil means surrounds a subject and generates a high-frequency magnetic field at the Larmor frequency and receives a spin resonance signal from the subject; a second high-frequency coil means disposed adjacent to the subject and inside the first high-frequency coil means for generating a motion signal; and an impedance of the second high-frequency coil means for determining the motion signal. An MRI tomography device comprising an impedance measuring means for measuring impedance. 2. The MRI tomography apparatus according to claim 1, wherein the second high-frequency coil means has a resonant frequency, and the resonant frequency is substantially higher than the proton resonant frequency. 3. The first high frequency coil means generates a first magnetic field, the second high frequency coil means generates a second magnetic field, and the second magnetic field extends substantially perpendicular to the first magnetic field. Claims 1 or 2, characterized in that the device is arranged so as to
MRI tomography device. 4. The second high-frequency coil means comprises a filter device, and the second high-frequency coil means causes the first high-frequency coil means to
The MRI tomography apparatus according to any one of claims 1 to 3, characterized in that it consumes only a small amount of energy. 5. The MRI tomography apparatus according to any one of claims 1 to 4, wherein the second high-frequency coil means is rigidly connected to the first high-frequency coil means. 6. Any one of claims 1 to 4, characterized in that the second high-frequency coil means is disposed so as to be slidable within one plane with respect to the first high-frequency coil means. MRI tomography device described in . 7. Claim 1, characterized in that the second high-frequency coil means is arranged so that it can be positioned independently with respect to the first high-frequency coil means.
MRI tomography apparatus according to any one of Items 1 to 4. 8. Claims 1 to 8, characterized in that the impedance measuring means measures the impedance of the second high-frequency coil means at a frequency slightly shifted from the resonant frequency of the second high-frequency coil means. MRI tomography device according to any one of clause 7. 9. Claims 1 to 7, characterized in that the impedance measuring means measures impedance at a frequency different from the spin resonance frequency measured by the first high-frequency coil means. MRI tomography device according to any one of the above.