Thermostabilization of antenna array for magnetic resonance tomography

A device for thermal stabilization of a first electrical characteristic of an antenna array of a magnetic resonance tomograph includes a heat exchanger configured for thermal coupling of a component of the device to a heat source. The device also includes a temperature-dependent second electrical characteristic. In a predefined connection to the antenna array, the temperature-dependent second electrical characteristic is configured to compensate for an effect of a temperature-dependent change on the first electrical characteristic of the antenna array in a predetermined temperature range.

RELATED APPLICATIONS

This application claims the benefit of German Patent Application No. DE 102013215918.8, filed Aug. 12, 2013. The entire contents of the priority document are hereby incorporated herein by reference.

TECHNICAL FIELD

The present teachings relate generally to devices for thermal stabilization of a first electrical characteristic of an antenna array of a magnetic resonance tomograph. The present teachings further relate to antenna systems and magnetic resonance tomographs including such devices.

BACKGROUND

Magnetic resonance measurements involve observing the interaction of magnetic moments of atomic nuclei (e.g., nuclear spins) with an external magnetic field.

When excited by an external alternating electromagnetic field around the axis of magnetic field orientation, nuclear spins align themselves in the external agnetic field and precess at a Larmor frequency that depends on the value of the magnetic moment of the nucleus and the external magnetic field. The atomic nuclei then generate an electromagnetic alternating field at the Larmor frequency.

The external alternating electromagnetic field used to excite the nuclear spins is projected into a sample or into a patient via one or a plurality of antenna arrays. One antenna array is a body coil that encircles the patient or the sample. However, local coils that are disposed directly on the patient or on the sample may be used. The electromagnetic field generated by the atomic nuclei is likewise received by the antenna arrays. The same antenna array may receive the signal that has been generated. Alternatively, the nuclear spins may be generated with one type of antenna and the electromagnetic alternating field generated by the atomic nuclei may be received using a different type of antenna.

The efficiency of projection and the sensitivity of reception are dependent on a plurality of characteristics of the antenna array (e.g., the electrical characteristics thereof). Characteristics of the antenna array may be the resonant frequency or the impedance. These characteristics of the antenna array are also dependent on the temperature of the antenna array and the components thereof. Thus, the inductance of a coil or the capacitance of a capacitor may be changed as a result of thermal expansion.

In the construction of the antenna arrays, the elements used may have a temperature coefficient that is equal to or close to zero for the mechanical or electrical characteristics.

The consequences of changes are also offset by control mechanisms. For example, lower reception sensitivity and/or transmission efficiency when the resonant frequency has changed may be compensated for by greater transmitting speed or input amplification.

However, the electrical characteristics may not always be kept constant using components having a low temperature coefficient because the characteristics of the antenna array also depend on the environment.

A body coil may be provided on a cylindrical element that is disposed concentrically between the patient or the sample and the gradient coils. The gradient coils do not prevent alternating electromagnetic fields from being beamed down onto the patient. In order to reduce external interactions with the gradient coils (e.g., to prevent irradiation and absorption of high-frequency energy in the gradient coils), a high-frequency shield may be disposed on the inside of a supporting base for the gradient coils. The shield extends between the gradient coils and the body coil. The body coil and the shield interact. For example, facing metal surfaces of the body coil and the shield effect a capacitive coupling. The electromagnetic waves that are transmitted by the antenna array generate eddy currents in the shield. Since the distance between the body coil and the shield changes if the gradient coil, together with the supporting base and the shield located thereon, becomes hotter, the electrical characteristics of the body coil change. The electrical characteristics change even if the body coil were to have a constant temperature or were configured with a temperature coefficient equal to zero.

SUMMARY AND DESCRIPTION

The present embodiments may obviate one or more of the drawbacks or limitations in the related art. For example, in some embodiments, a magnetic resonance tomograph is provided wherein the temperature-related effects on the antenna array are reduced.

A device in accordance with the present teachings configured for the thermal stabilization of a first electrical characteristic of an antenna array of a magnetic resonance tomograph includes a heat exchanger for thermal coupling of a component of the device to a heat source or to the antenna array. All elements that allow heat transfer between the heat source and a component of the device may be used as heat exchangers in accordance with the present teachings. Examples include fasteners, sealing compounds, structural elements that the component is embedded into, and cooling devices pertaining to the heat source (e.g., configured for exchanging heat with the component). The device further includes a temperature-dependent second electrical characteristic that is configured, in a predefined connection to the antenna array, to compensate for an effect of a temperature-dependent change caused by the heat source on the first electrical characteristic of the antenna array in a predefined temperature range. The second electrical characteristic of a circuit that includes a plurality of components may be a temperature-dependent second electrical characteristic in accordance with the present teachings.

In a circuit having a predetermined antenna array, a device in accordance with the present teachings may be in thermal coupling with the antenna array itself or an object that causes a temperature-related change in the first electrical characteristic of the antenna array. A device in accordance with the present teachings may compensate for a change in the first electrical characteristic through a temperature-related change in the second electrical characteristic through connection to the antenna array. The first electrical characteristic of a system that includes the antenna array and a device in accordance with the present teachings remains substantially unchanged in a predefined temperature range. The phrase “substantially unchanged” as used in this context refers to the value of the first electrical characteristic changing only slightly (e.g., by a maximum of 1%, 2%, 5% or 10%). A temperature range in accordance with the present teachings may encompass temperatures between 20° C. and 30° C., 20° C. and 40° C., and 15° C. and 50° C.

An antenna system in accordance with the present teachings has a device in accordance with the present teachings and an antenna array. The device is thermally coupled to the antenna array via the heat exchanger. The device is electrically connected to the antenna array, such that the antenna system has a substantially unchanged first electrical characteristic in the predefined temperature range.

By thermal coupling between the antenna array and the device, the antenna system in accordance with the present teachings may have properties for the first electrical characteristic that are substantially independent of the temperature in a predefined temperature range.

A magnetic resonance tomograph in accordance with the present teachings includes a device, an antenna array, and a heat source. The heat source is in a cause-and-effect relationship with a temperature-dependent change in the first electrical characteristic. The magnetic resonance tomograph includes a first thermal coupling between the device and the heat source via the heat exchanger. The device is electrically connected to the antenna array, such that an effect of the temperature-dependent change on the first electrical characteristic of the antenna array is substantially offset in a predefined temperature range.

If the change in the first electrical characteristic of the antenna array is not directly dependent on the temperature of the antenna array itself, but rather is indirectly caused by temperature-dependent physical effects on the antenna array, the first electrical characteristic of a system that includes the antenna array and the device may be kept unchanged in a predefined temperature range using a magnetic resonance tomograph and device in accordance with the present teachings.

In some embodiments, a device in accordance with the present teachings includes only passive components. As used herein, the phrase “passive components” refers to components that do not require any further power supply or control signals for utilization and that function apart from a signal to be processed. For example, passive components may be resistors, coils, and capacitors.

Since a device in accordance with the present teachings may have only passive components, the device may be inserted into existing systems.

In some embodiments, the first electrical characteristic is a resonant frequency of the antenna array.

In a resonant antenna array, the efficiency of transmission and the sensitivity of reception of an electromagnetic wave are dependent on the resonant frequency of the antenna array being consistent with the frequency of the electromagnetic wave. Since a device in accordance with the present teachings is configured to compensate for a temperature-related change in the resonant frequency of the antenna array in a predefined temperature range, the device may be used to reduce or eliminate this dependency in a predefined connection to the antenna array.

In some embodiments, the first electrical characteristic is an impedance of the antenna array.

The efficiency of transmission and the sensitivity of the reception of an electromagnetic wave substantially depend on the impedance of the antenna array being aligned with the impedance of the power supply. If the two impedances are not consistent, losses occur as a result of the reflection of the electric signals at the point where the antenna array interfaces with the power supply. Since a device in accordance with the present teachings is configured to compensate for a temperature-related change in the impedance of the antenna array in a predefined temperature range, the device may be used to reduce or eliminate this dependency when in a predefined connection to the antenna array.

In some embodiments, the second electrical characteristic is a capacitance of the device.

A variable capacitance may be used to compensate for different first electrical characteristics in the temperature properties of the circuits by having different circuits. For example, capacitances may be achieved in a small spatial area.

In some embodiments, the heat exchanger is a metal contact surface. A metal contact surface may have a good thermal conductivity.

In some embodiments, the heat exchanger is a conduit for a cooling medium. The conduit may be used to provide a thermal coupling with a device in accordance with the present teachings that does not require direct vicinity and that allows electrical isolation from a heat source.

In some embodiments, the heat source is a gradient coil. A thermal coupling with the gradient coil may be used to compensate for changes in the first electrical characteristic of the antenna array that have been caused by thermal effects (e.g., expansion of the gradient coil) even if the antenna array itself does not become hot.

In some embodiments, the magnetic resonance tomograph further includes a coolant circuit to cool the heat source. The heat exchanger is in thermal contact with the coolant circuit. The coolant circuit may be used to provide a thermal coupling of the heat source with a device in accordance with the present teachings that does not require direct vicinity and that allows electrical isolation from the heat source.

In some embodiments, the first thermal coupling has a first time constant for a first heat transfer between the heat source and the device, and the second thermal coupling between the antenna array and the heat source has a second time constant. The first thermal coupling is configured such that the antenna array has a substantially unchanged first electrical characteristic for a predefined operating profile of the magnetic resonance tomograph.

The heat transfer between the heat source and the device, and the change in the first electrical characteristic that is dependent on a temperature change in the heat source, may have different characteristics. Since the time constants are aligned with one another by design features (e.g., length of the thermal conduction pathways, flow speed of the cooling medium, thermal capacities or electrical characteristics of components), compensation for the first electrical characteristic may be achieved even where there is a dynamic temperature progression in an operating profile of the magnetic resonance tomograph.

DETAILED DESCRIPTION

The magnetic resonance tomograph1includes a magnet unit10with a field magnet11. The field magnet11generates a static magnetic field B0to direct nuclear spins of samples or of a patient40in a sample volume. The sample volume is disposed in a duct16that extends in a longitudinal direction2through the magnet unit10. The field magnet11may be a superconducting magnet that may provide magnetic fields with a magnetic flow density of up to 3T and, in some machines, even higher. For lower field intensities, permanent magnets or electromagnets with normally conducting coils may be used.

Furthermore, the magnet unit10includes gradient coils12that are configured to superimpose variable magnetic fields in three spatial directions on the magnetic field B0in order to spatially differentiate the imaging zones that have been captured in the sample volume. The gradient coils12may be coils of normally conducting wires that, in the sample volume, may generate fields that are orthogonal to one another.

On the inside of the gradient coils12and disposed towards the sample volume16, there is a shield13. The shield13has conductivity and substantially prevents propagation of high-frequency electromagnetic waves between the gradient coils12and the area located within the shield13(e.g., waves that have a frequency range above 1 MHz, as do the waves used in magnetic resonance tomography). The shield13is disposed on the gradient coils or on a common supporting base.

The magnet unit10also includes a body coil14and local coils15. Both the body coil14and the local coils15may be referred to as antenna arrays14,15in the description that follows. Both the body coil14and the local coils15may emit a high-frequency alternating magnetic field into the surrounding area. The body coil14is used inter alia as a transmission coil to generate across a large volume a homogeneous electromagnetic excitation field.

The local coils15may be arranged as a two-dimensional or three-dimensional matrix and cover parts of the body of the patient40. The local coils15are used inter alia as transmission coils in order to project, in each case, electromagnetic waves into a substantially spatially limited volume of the body. The local coils15may, for example, be circular or polygonal coils that partially overlap one another. The fields of adjacent coils may be partially superimposed on one another (e.g., some in the same direction and some in opposite directions), such that adjacent coils substantially do not interact with one another. The overlapping arrangement of the transmission coils15may be used to project an alternating electromagnetic field in the entire area to be examined that is covered by the coils.

A magnetic resonance signal that is generated by the electromagnetic field of the body coil14or the local coils15and the static magnetic field B0in the patient may either be picked up again by the local coils15or by the separate body coil14. The separate body coil14may receive signals from the entire area that is being investigated.

A control unit20supplies the magnet unit10with the various signals for the gradient coils12and the body coil14or the local coils15. The control unit20evaluates the signals that have been received.

The control unit20includes a gradient driver21that is configured to supply the gradient coils12via cables with variable currents that are time-coordinated. The control unit is further configured to provide the desired gradient fields in the sample volume.

The control unit20includes a transmitting and receiving unit22that is configured to generate a high-frequency pulse with a predefined time progression, amplitude, phase, and spectral power distribution for an antenna array14,15in order to generate a magnetic resonance of the nuclear spins in the patient40, thereby creating pulse outputs in the kilowatt range.

The transmitting and receiving unit22is further configured to evaluate (e.g., for amplitude and phase) high-frequency signals that have been received from the body coil14or one or a plurality of local coils15and supplied via a signal circuit33to the transmitting and receiving unit22. These signals may be high-frequency signals that transmit nuclear spins in the patient40in response to excitation by a high-frequency pulse in the magnetic field B0or in a magnetic field resulting from a superimposition of B0and gradient fields.

The control unit20further includes a control23that is configured to carry out the time coordination of the activities of the gradient driver21and the transmitting and receiving unit22in order to capture images using magnetic resonance tomography. The control23is connected to the other gradient driver21and the transmitting and receiving unit22via a signal bus25in a signal exchange. The control23is configured to accept and process signals from inside the patient40that have been evaluated by the transmitting and receiving unit22, or to provide the gradient driver21and the transmitting and receiving unit22with pulse and signal forms and to coordinate the pulse and signal forms with respect to time.

The patient40is placed on a patient table30as are used in magnetic resonance tomography. The patient table30includes a first supporting strut36that is arranged beneath a first end31of the patient table30. To maintain the patient table30in a horizontal position, the supporting strut36may have a foot that extends along the patient table30. In order to move the patient table30, the foot may also include a moving element (e.g., rollers). Apart from the supporting strut36at the first end31, there are no structural components between the floor and the patient table. As a result, the patient table may be slid up as far as the first end31into the duct16in the field magnet11.FIG. 1shows linear rail systems34that moveably connect the supporting strut36to the patient table30, such that the patient table30may move in a longitudinal direction2. The linear rail system includes a drive37whereby the patient table30may be moved in a longitudinal direction2. The operation is controlled either by an operator or by the control23. As a result, areas of the patient's body that cover a greater expanse than the sample volume may be investigated in the duct16.

The magnet unit10includes a cooling system70that supplies a cooling medium via cooling medium lines71to the gradient coils for cooling and then returns the cooling medium via the cooling lines71to the cooling system70. The cooling medium releases heat energy to the cooling system70via a heat exchanger72.

FIG. 1shows an exemplary first device50and an exemplary second device60in accordance with the present teachings. The first device50and the second device60may be used for the thermal stabilization of a first electrical characteristic of an antenna array14,15. The first device50and the second device60shown inFIG. 1are two exemplary devices that may be connected to a magnetic resonance tomograph1in two ways. However, other kinds of connections may be used. For example, the simultaneous representation of the first device50and the second device60is only one example of a connection.

The device50is connected via the signal circuit33to the transmitting and receiving unit22and to the cooling system70via cooling medium conduits71. The device50is arranged in the cooling medium conduit such that the cooling medium heated by the gradient coil12flows through the device50and has a temperature that indicates a measurement of the temperature of the gradient coil12and the shield13. In some embodiments, the temperature of the cooling medium is equal to the temperature of the gradient coils12. In other embodiments, a temperature difference in the cooling medium in the device50is proportional to a temperature difference between the gradient coil12and the ambient temperature. Other correlations may also be used.

Further details of the device50are depicted inFIG. 2.

As shown inFIG. 3, the device60is in direct thermal contact with a local coil15. The device60is connected to the signal circuit33, such that an outgoing or incoming signal from or to the local coil15passes through the device60.

The device60may be arranged in direct thermal contact with the body coil15and inserted into the signal circuit33between the transmitting and receiving unit22and the body coil14.

A plurality of local coils15may be used in a magnetic resonance tomograph in accordance with the present teachings. Each individual coil may be provided with a device60.

In some embodiments, separate devices (e.g., first device50and second device60) may be provided in each case for transmitting and receiving signals. Transmission may occur using the body coil14, and reception may occur using local coils15or vice versa. The first device50and the second device60may be assigned in each case to the antenna arrays14,15.

FIG. 2shows a schematic diagram of an embodiment of a device50in accordance with the present teachings.

The device50includes a conduit51in a heat exchanger52. Cooling medium conduits71of a cooling system70are connectable to the heat exchanger52. The cooling medium may circulate through the conduit51. The heat exchanger52for thermal coupling may acquire the temperature of the cooling medium.

The device50further includes a first capacitor53and a third capacitor56that are in thermal contact with the heat exchanger52and that acquire the temperature of the heat exchanger52. As a second electrical characteristic, the first capacitor53and the third capacitor56have a capacity that is dependent on the temperature. Together with the second capacitor55and the coil54, the capacitors53,55,56form an adaptor box that adapts the impedance of the body coil14in the signal circuit33to the transmitting and receiving unit22. The dotted circuit inFIG. 2shows an equivalent circuit diagram for the body coil14with an antenna impedance56in series with an antenna capacitor57.

Through appropriate selection of the temperature coefficient of the first capacitor53and the third capacitor56, compensation for the change in the first electrical characteristic of the body coil14caused, for example, by an expansion of the shield13may be achieved.

In some embodiments, the body coil14includes an antenna impedance57of 60 Ohms and an antenna capacity58of 20 pF. The power cable has an impedance of 50 Ohms. The first capacitor53is configured with a capacity of 10 pF, the second capacitor55has a capacity of 26.6 pF, the third capacitor56has a capacity of 10 pF, and the coil54has an inductance of 265 nH. The coil54and the second capacitor55have a temperature gradient that is substantially equal to zero and/or the coil54and the second capacitor55are maintained at a constant temperature by, for example, being thermally isolated from the heat exchanger52.

In an exemplary magnetic resonance tomograph1, the gradient coils12heat up by 30 degrees centigrade when the gradient coils12are in operation. As a result, the temperature of the shield rises by 25 degrees centigrade. The body coil14itself heats up by 20 degrees centigrade. The resonant frequency of the body coil drops by 250 kHz. The drop corresponds to an increase in the antenna capacity58in the equivalent circuit diagram for the body coil14from 20 pF to 20.1 pF.

To compensate for the change, the capacity of the first capacitor53rises as a result of being heated up to 10.18 pF, and the capacity of the third capacitor56drops to 9.97 pF. If the temperature of the cooling water is equal to the temperature of the capacitors53,56, there is a positive temperature coefficient for the first capacitor53of 6 *10−4l/K, and a temperature coefficient for the third capacitor56of −1*10−4l/K.

The capacitors are made, for example, using dielectrics with a low dielectricity constant (e.g., LDC capacitors). For example, special steatites or earthing elements containing rutile (TiO2) are used. The dielectrics of this type are used in the manufacture of temperature-coefficient capacitors. By using additives, the negative temperature coefficient of TiO2 of −800*10-6/K may be moved up to zero and even into the positive range. With different additives, a further move into the negative range is achieved. In this way, materials with a temperature coefficient of +100, ±0, −33, −75, −150, −470 and −1500*10-6/K are obtained. These materials are known, respectively, as P100, NP0, N33, N075, N150, N470, and N 1500.

Higher positive temperature coefficients may be achieved, for example, by using the dielectric barium titanate BaTiO3as a base material. Barium titanate has a relative dielectricity constant (DC) of several thousand at a temperature of about 120° C. Above and below this temperature point (e.g., the “Curie point”), the DC drops with a value of 1500 being produced at a temperature of 20° C.

FIG. 3shows a further embodiment of a device60in accordance with the present teachings. The embodiment shown inFIG. 3differs from the embodiment shown inFIG. 2in that the heat exchanger does not include a cooling medium conduit. Instead, heat exchange occurs through direct contact. InFIG. 3, elements identical to those shown inFIG. 2are denoted by the same reference signs.

InFIG. 3, the heat exchanger52is in direct contact with the coil. The coil may be a body coil14or a local coil15. The heat exchanger52may be disposed directly on the gradient coils12.

The contact surface for the coil14,15may be a flat surface that is made of metal and abuts onto a corresponding surface of the coil14,15. Other complementary surfaces on the coil and the heat exchanger configured to come into full surface contact with each other may be used instead of a flat surface. In some embodiments, the heat exchanger52may include a bolt that is screwed into a corresponding thread on the coil.

In some embodiments, the heat exchanger52may be formed by a component of the coil14,15or gradient coil12or of the shield13thereof. The capacitor or capacitors53,56may be embedded, for example, in an epoxide resin that forms the structure of the gradient coil12or the coils14,15.

Because the coil54and the second capacitor55are in the direct vicinity of the heat source, the coil54and the second capacitor55may not be prevented from heating up even with thermal insulation. In the embodiment shown inFIG. 3, the elements54,55may have a temperature coefficient that is substantially zero. The elements54,55interacting in the circuitry provided may compensate reciprocally for the thermal changes in their electrical values.

In some embodiments, a device in accordance with the present teachings may be configured as a component of the coils14,15. For example, the device may be completely embedded in a casting resin that forms a coil body of the body coil14or of the local coil15.