Method and apparatus for mobile MRI with permanent magnets

Abstract

Methods and apparatus are provided to implement a low-field MRI system that can be used in hospital departments and ICUs and move therein for point-of-care. The low field MRI system is implemented by a number of features. The moving magnet provides a stable B0 field generated by the permanent material. In order to increase the stability of the B0 field, a monitoring coil and a shim coil are operated in a combined manner. In addition, several non-50 ohm circuit methods for use in moving MRI systems are provided. Transmit circuits are provided that include different bleeder and detune circuits. In a receive channel, an array of receive coils using parallel tuned circuits is provided to implement multi-channel reception. In addition, decoupling and far-away detuning strategies are provided.

Description

Technical Field

[0001] The present invention generally relates to low field mobile methods and apparatus for performing diagnostic imaging. More particularly, the present invention relates to brain magnetic resonance imaging apparatus. Background Art

[0002] Magnetic resonance imaging (MRI) technology has been well used for in vivo imaging in the medical field. When human tissue is in a static magnetic field B0, the nuclear spins in the tissue are polarized and have a net alignment moment Mz. If the tissue is subjected to a radio frequency (RF) field B1 at the Larmor frequency in the xy plane, then Mz can be tilted into the xy plane to produce a net transverse magnetic moment Mt. Then, gradient fields (Gx, Gy, and Gz) can be used to select nuclear spins in certain areas, and spatial encoding can be performed to generate spatial information. Signals are emitted by the excited nuclear spins, and this signal can be detected and processed using one or more RF coils to produce an image.

[0003] To address the low sensitivity in inductive detection of weakly polarized nuclear spins, most clinical MRI scanners employ superconducting magnets to generate high magnetic fields (e.g., operating at 1.5 T or 3 T or even 7 T). For optimal performance, the magnet should typically have an inhomogeneity of about one part per million (ppm) over the imaging volume. While increased field strength clearly provides the advantage of a higher signal-to-noise ratio (SNR), it also brings several practical limitations, including increased magnet size and weight, increased specific absorption rate (SAR, which scales with B02), increased inhomogeneities in both the static field B0 and the applied RF field B1, increased likelihood of heating of the implanted device, and increased image artifacts around metallic devices.

[0004] The use of low-field MRI devices that utilize permanent magnets can solve these problems. In addition, the use of permanent magnets can make the MRI device mobile and potentially useful for point-of-care, such as bedside diagnosis or a doctor's office.

[0005] Therefore, there is a need for systems and methods that address the above limitations.The following disclosure encompasses an improved mobile MRI system with a stable B0 field that employs a more efficient approach to pulse transmission and signal reception. Summary of the invention

[0006] Embodiments of systems and methods for implementing a mobile MRI device are disclosed, which can be used in clinical diagnosis and human research, such as that required for brain MRI. Certain embodiments disclosed herein include methods for implementing simpler and more efficient MRI circuits for transmission and reception. Certain embodiments disclosed herein also include methods for stabilizing the B0 field.

[0007] In an embodiment, a method for stabilizing the magnetic field of a mobile MRI device using a permanent magnet is provided. The method includes: providing a main static magnetic field by a permanent magnet of the mobile MRI device; providing a radio frequency pulse by a frequency measurement module of the mobile MRI device to induce a free induction decay (FID) signal; monitoring the phase of the FID signal by the frequency measurement module to determine the frequency of the main static magnetic field; and adjusting the frequency of the main static magnetic field by a compensation module.

[0008] In an embodiment, a system for transmitting RF fields in a mobile MRI apparatus including a transmit channel is provided. The transmit channel includes a class D power amplifier, an anti-parallel diode, one or more circuits configured for detuning or bleed, and at least one of a tuned RF coil and an untuned RF coil.

[0009] In another embodiment, a mobile MRI device is provided. The mobile MRI device includes a permanent magnet configured to provide a main static magnetic field; and a receiving channel. The receiving channel includes a parallel tuned probe, a preamplifier, at least one detuning circuit, an impedance converter, and a signal acquisition device. BRIEF DESCRIPTION OF THE DRAWINGS

[0010] Figure 1 1 is a general diagram showing a mobile MRI system having the functions of field stabilization and imaging according to an embodiment of the present invention.

[0011] Figure 2A An MRI system with a permanent magnet and an array receiving coil according to an embodiment of the present invention is shown.

[0012] Figure 2B Aspects of a receive coil employing a geometric decoupling strategy according to an embodiment of the present invention are shown.

[0013] Figure 3A A mobile MRI system including its mechanical assembly according to an embodiment of the invention is shown.

[0014] Figure 3B A mobile MRI system including its mechanical assembly according to an embodiment of the invention is shown.

[0015] Figure 4 FIG. 4 is a schematic diagram showing the hardware of a field monitoring channel according to an embodiment of the present invention.

[0016] Figure 5 FIG. 4 is a flow chart illustrating a process of monitoring and compensating for a magnetic field according to an embodiment of the present invention.

[0017] Figure 6 Schematic diagram of a transmit channel using a tuning coil in combination with a detuning and bleeder circuit according to an embodiment of the present invention.

[0018] Figure 7 A schematic diagram showing hardware for a Class D RF power amplifier according to an example embodiment.

[0019] Figure 8 is a schematic diagram of a transmitting coil employing a parallel tuning circuit used with a bleeder and detune circuit according to an embodiment of the present invention.

[0020] Fig. 9 is a schematic diagram of a transmitting coil employing a series tuning circuit used with a bleeder and detuning circuit according to an embodiment of the present invention.

[0021] Fig.10 FIG. 1 is a schematic diagram of a transmission channel using an untuned coil and combined with a bleeder circuit according to an embodiment of the present invention.

[0022] Fig.11 is a schematic diagram of a transmitting coil employing an untuned circuit for use with a bleeder circuit according to an embodiment of the present invention.

[0023] Fig.12 FIG. 4 is a general schematic diagram of a receiving channel with a detuning circuit according to an embodiment of the present invention.

[0024] Fig.13 FIG. 4 is a schematic diagram showing a detuning and preamplifier decoupling strategy in a receiving channel according to an embodiment of the present invention.

[0025] Fig.14 is a graph of induced current measurement using a sniffer loop according to an embodiment of the present invention. DETAILED DESCRIPTION

[0026] The present disclosure provides systems, methods, and devices for improving mobile MRI systems. Mobile MRI systems can be easily deployed for point-of-care, for example, in an ICU or a doctor's office. The development of mobile MRI devices can have several significant advantages.

[0027] One of the advantages of using a low-field device is that the Lorentz force is reduced, and the reduction of the Lorentz force usually greatly reduces the acoustic noise level, which is very desirable for patient research. Another advantage is that the longitudinal relaxation time (T1) of the tissue decreases with the decrease of the magnetic field B0, while the transverse relaxation time (T2* / T2) increases with the decrease of the magnetic field B0. By using such a low-field mobile MRI device, the shorter T1 and longer T2 of the tissue may be beneficial for acquisition with reduced polarization time and long echo chains. Another advantage of using a low field is that it is possible to use shorter RF pulses (because no SAR limitation is expected), as well as more uniform contrast due to increased B1 uniformity. These options can open up the potential for new technology applications, such as fast spin echo (FSE) that may have problems at high field strengths. Therefore, if the image quality is acceptable for clinical applications, the mobile MRI device has many advantages.

[0028] The uniformity and stability of the static magnetic field B0 are crucial for mobile MRI systems using permanent magnets. The size of the homogeneous region of the magnetic field is defined as the spherical volume diameter (DSV). Magnet homogeneity is directly related to image quality and various artifacts (e.g., blurring). Another important factor is that inhomogeneities in the B0 field can reduce T2*, which can be a disadvantage for some sequences such as gradient echo sequences.

[0029] When using permanent magnets, it is also important to ensure temperature stability of the B0 field. When the room temperature changes by even one degree (which may be normal in a centrally air-conditioned hospital room), the B0 field may change or the uniformity may deteriorate.

[0030] Class D type RF power amplifiers have been used in low frequency NMR / MRI applications and can achieve high efficiency. This type of RF amplifier also uses parallel tuned circuits, non-resonant circuits and even series tuned circuits to implement the transmission coil. These circuits can be relatively easy to build and the resonant frequency can also be changed by a switched capacitor when necessary. On the receiving side, by using an array-based receiving coil, impedance matching and tuning circuits (50 ohms) can be used and combined with preamplifier decoupling technology. The parallel tuned circuit can have the advantages of an impedance matching circuit in terms of high signal gain. In addition, a large value inductor with high Q for use in the preamplifier decoupling circuit may be required in a low field MRI device. The preamplifier decoupling method proposed in the present disclosure can use fewer lumped components. The parallel tuned circuit on the coil can also have the benefit of using fewer lumped components to reduce signal loss.

[0031] The present disclosure includes a method for stabilizing a magnetic field by using a frequency measurement module and a compensation module. For MRI systems based on permanent magnets, the ability to stabilize the B0 field is necessary to obtain good image quality. The advantages of embodiments of the present disclosure are not limited to mobile MRI systems and can be used to improve field stability in any MRI system.

[0032] The present disclosure includes a method for simplifying the transmit channel of a mobile MRI system using tuned circuits and untuned circuits. In addition, corresponding methods for bleed and detuning are presented herein. The ability of such circuits can reduce the effect of the transmit coil when the transmit coil is not in use.

[0033] The present disclosure includes a method for a mobile MRI system using a combination of parallel tuned array coils and preamplifier decoupling techniques. In addition, a detuning method using off-coil circuits is used to deactivate the receive array coils when they are not in use.

[0034] Although described with respect to mobile MRI systems, the features and embodiments discussed herein are not limited to use with such mobile MRI systems.

[0035] Figure 1 An example embodiment of a mobile MRI device 10 based on a permanent magnet configured to provide a main static magnetic field B0 is shown. The system includes several modules for implementing the following two functions: stabilizing the static magnetic field and performing imaging. The spectrometer 101, which serves as a key component of the mobile MRI device 10, is used to connect these modules together and interface with a computer configured to interpret the results. In this example system, a permanent magnet 102 is used to generate a uniform static magnetic field stabilized by a frequency measurement module 103 and a compensation module 104 (collectively referred to as a monitoring subsystem). A computer 100 is connected to the spectrometer 101 and is configured to monitor the magnetic field changes measured by the frequency measurement module 103. The field change measurement of the magnetic field is transmitted to the computer 100 and analyzed. The computer 100 then controls the spectrometer 101 to adjust the magnetic field by the compensation module.

[0036] like Figure 1As shown in , MRI sequence information is transmitted from a computer 100 to a spectrometer 101, which controls a gradient module 106, a transmission module 105, and a receiving module 107, respectively. The transmission module 105 uses a short duration RF field to tilt the nuclear spin magnetization or focus it together. The gradient module 106, which may include three gradient amplifiers, is configured to select spins in a region of interest or perform spatial encoding; and the receiving module 107 is configured to measure the magnetic field generated by the induced spins and amplify these signals. By combining the functions of the transmission module 105, the gradient module 106, and the receiving module 107, an MRI imaging function can be implemented. After being acquired by the spectrometer 101, the raw imaging data can be transmitted to the computer 100. The raw imaging data can be uploaded to a cloud computing service center and / or processed by the computer 100. After processing, the raw image data can be transmitted back to the computer 100 and displayed, or displayed only when the processing is performed locally. The processing program may include a denoising algorithm not discussed herein.

[0037] Figure 2A An example embodiment of one aspect of a mobile MRI system 10 or device is shown. The magnet 102 includes an iron yoke 200 and magnetic elements 201 (i.e., magnet blocks, passive shimming materials, and shimming rings). Each magnet block generates a static magnetic field. The passive shimming material may include small pieces of permanent magnets. It may also include small pieces of metal. The iron yoke 200 using a high magnetic permeability material can provide a magnetic path for the static field and also reduce the leakage of magnetic flux. The thickness of the iron yoke can be selected to avoid magnetic saturation. For example, an iron yoke 200 with a thickness of 35 mm can be used with a 0.1 T magnet. The shimming rings, which are also made of high magnetic permeability material, can constrain the magnetic field and ensure that the main magnetic field B0 is more uniform at a certain DSV. In an embodiment, the magnet 102 may have a main magnetic field in a vertical direction, such as Figure 2A As shown in .

[0038] The magnetic element 201 may include an X / Y / Z gradient coil. The three gradient coils may be connected to the gradient module 106. According to some embodiments, a second order gradient coil (i.e., a Z2 gradient coil) may also be included in the magnetic element 201. The spectrometer 101 in communication with the gradient module 106 may control a gradient power amplifier to amplify the gradient wave and deliver it to the corresponding gradient coil. The gradient coil may use an offset current and have a shimming function to make the static magnetic field more uniform. The offset current on the gradient coil may vary with magnetic field fluctuations. A pipe for water cooling may be provided around the gradient coil. The water may be pumped with a small DC motor. This may ensure that the heat from the gradient coil does not affect the permanent magnet. In an embodiment, a heat insulating material may be used between the permanent magnet and the gradient coil. An interface 202 is provided for connecting the gradient coil and the water pipe.

[0039] The compensation module 104 may include a shim amplifier and a shim coil 204. The shim coil 204 may act as a compensation coil to stabilize the static main magnetic field provided by the permanent magnet 102. The spectrometer 101 controls the output current of the shim amplifier connected to the shim coil 204, which supplies DC power (<2W) to the compensation coil 204. In some embodiments, the shim coil 204 may use several coils to stabilize the static field. In an embodiment, the spectrometer 101 may control the shim coil 204 in conjunction with the shimming function of the gradient coil of the magnetic element 201 to ensure the stability and uniformity of the magnetic field B0.

[0040] In an embodiment, the mobile MRI device 10 may include a transmit coil connected to or associated with the transmit module 105. To generate a uniform B1 field, the transmit coil may be a solenoid coil or a saddle coil disposed outside the receive coil 203.

[0041] The receiving coil 203 may be an array coil comprising 9 elements, including, for example, one conical solenoidal coil, four figure-8 coils, and four surface coils. Other suitable arrangements of the receiving coil 203 may also be selected. The coils in the array of the receiving coil 203 may be arranged so that the RF field in the sensitive region is perpendicular to the B0 field in the vertical direction. Figure 2B An example of detailed geometry of a surface coil 300 of a receive coil 203 and a figure-eight coil 301 of the receive coil 203 is shown. The surface coil 300 and the figure-eight coil 301 may have more than one turn. There may be overlap between adjacent coils of the array of receive coils 203 to reduce coupling (e.g., geometric decoupling). In addition to geometric decoupling, preamplifier decoupling methods may also be used in the array of receive coils 203, as further described below.

[0042] Figure 3A The mechanical assembly of the mobile MRI system 10 on wheels is shown. Figure 3A The mobile MRI system 10 is shown in the MRI measurement state, wherein the magnet door 410 is open. To avoid being attracted to the magnetic metal parts of the magnet, the magnet door 410 can be closed when the device is not in use, such as Figure 3B As shown in FIG. 414 ( Figure 3B ) and support plate 416 can also be folded upward. Magnet door 410 can be aluminum, copper or other suitable materials. Cover portion 414 and support plate 416 can be plastic or other non-magnetic materials. Non-magnetic steel structure 412 can be manufactured by welding process to be arranged on six wheels 413. One or more of wheels 413 can be power-assisted electric wheels. More or fewer wheels 413 can be used. Handle 411 can be used to move MRI system 10. Handle 411 can be folded during MRI measurement. Handle 411 can include control members such as triggers, buttons, etc. to control the power-assisted electric wheels.

[0043] Figure 4 FIG. 1 is a schematic diagram of an exemplary hardware of the frequency measurement module 103. The frequency measurement module 103 is a small NMR system controlled by a computer 100 and a spectrometer 101. M RF pulses 400 are provided by the spectrometer 101 and delivered to a power amplifier 401, which sends RF power (<10W) to a tuned RF monitoring coil 404 through a pair of diodes 402. The subscript M here indicates the monitoring coil 404. The RF monitoring coil 404 can be a solenoid coil or a saddle coil, and can have high sensitivity inside. In order to avoid the effect of human MRI images during field measurement, a small sample without hydrogen can be inserted into the RF monitoring coil 404 for measuring / monitoring the main static magnetic field B0. The sample can contain, for example, carbon 13 ( 13 C) or fluorine 19( 19 F), each of which has 1 H) have different gyromagnetic ratios. Therefore, the resonant frequency of the sample can be different from that of hydrogen. For example, hydrogen ( 1 H) and carbon 13 ( 13 C) have resonant frequencies of 42.58 MHz and 10.7 MHz respectively at 1 Tesla.

[0044] The detuning and bleeder circuit 403 may be used in the frequency measurement module 103. The detuning and bleeder circuit 403 is configured to deactivate the RF monitoring coil 404 and thereby reduce coupling with the transmit coil and receive coil 203 during MRI imaging. The detuning and bleeder circuit 403 may reduce the ring-down time of the RF pulses. Examples of the detuning and bleeder circuit 403 are provided in more detail below. A duplexer 405, which may be, for example, a switch based on a π circuit or a MOSFET, is also employed. The detuning and bleeder circuit 403 and the duplexer 405 may be in an off state when the RF monitoring coil 404 transmits pulses. When the RF monitoring coil 404 is used as a receive coil, the NMR signal received by the RF monitoring coil 404 passes through the duplexer 405 and is collected by the signal acquisition circuit 406, which is described below with respect to Fig.12 Describe in more detail.

[0045] As discussed above, the shim coils 204 and the frequency measurement module 103 may be combined to stabilize the magnetic field. Figure 51 is a flow chart showing a method for monitoring and compensating for a static magnetic field B0. At step 500, when the mobile MRI system 10 is powered on, the frequency measurement module 103 waits for a command to implement a test frequency. At step 501, the RF monitoring coil 404 of the frequency measurement module 103 is used to collect a free induction decay (FID) signal when the computer 100 provides an appropriate command. Collecting the FID signal may include: providing a radio frequency pulse by the frequency measurement module 103 of the mobile MRI device 10 to induce the FID signal; and monitoring or measuring the phase of the FID signal by the frequency measurement module 103 to determine the frequency of the main static magnetic field.

[0046] Fluctuations in the magnetic field B0 will cause fluctuations in the phase in the FID signal. At step 502, the frequency change in the magnetic field B0 can be evaluated based on the measurement of the phase Φ(t) of the FID signal. This can be illustrated using the expression Φ(t)=Φ0+ω*t, where Φ0 is the initial phase. ω(=γ*B) is the linear slope of the phase function. At step 503, this result can be compared with the initial frequency (or linear slope), and the fluctuation of the magnetic field (ΔB) can be obtained to determine whether the frequency has changed. If the frequency has not changed, the method can return to step 500 and wait for a command to make a new measurement. If the frequency has changed, the system can pass control to step 503.

[0047] At step 503, it is determined whether the frequency has been increased or decreased. If the frequency has been increased, the system is configured to decrease the current in the shim coil 204 at step 504. If the frequency has been decreased, the system is configured to increase the current in the shim coil 204. These steps may be repeated to maintain the stability of the magnetic field provided by the permanent magnet 102. Imaging may be performed after the magnetic field is stabilized.

[0048] Figure 6 A schematic diagram of a transmit channel 610 using a transmit coil 604 is shown, which also includes a detuning and bleeder circuit 603. The transmit channel 610 in a mobile MRI device will have an angular frequency ω 0 RF pulses 600 (eg, received from a spectrometer 101 as controlled by a computer 100) are delivered to an RF power amplifier 601. The amplified pulses pass through a diode 602 for application to an RF transmit coil 604, which generates an RF field that acts on the nuclear spins.

[0049] In an embodiment, a class AB power amplifier may be used in a high-field human MRI system. In an embodiment, a class D amplifier may be a first choice for a low-field NMR / MRI device. Compared to a linear amplifier (i.e., class AB), a class D amplifier may have better power efficiency and lighter weight. It may be low energy consuming. These features provide advantages for use in a mobile MRI system such as the mobile MRI device 10.

[0050] Figure 7 A schematic diagram of a voltage-mode class-D power amplifier 710 with a low impedance output according to an example embodiment is shown. Amplifier 710 is a switch-mode power amplifier including switches 702a, 702b, 703a, and 703b. Switches 702a, 702b, 703a, and 703b may be MOSFET-based switches. Switch 702 (702a and 702b) is controlled by signal 1, while switch 703 (703a and 703b) is managed by signal 2. Signal 1 and signal 2 ensure that switches 702 and 703 are not in a closed state at the same time. When the switches are closed, a high voltage 704 and a low voltage 705 are applied to two terminals of a load 706 (e.g., a transmitting coil 604), respectively. These switches may implement frequency modulation of a DC power supply 701, while a power supply controller 700 is used for amplitude modulation. Amplitude- and frequency-modulated RF power is applied to a load 706. According to some embodiments, a filter may be included to remove output harmonics of the class-D power amplifier.

[0051] Figure 7 The load 706 in the embodiment may be a tuning coil having a parallel circuit or a series circuit, such as Figure 8 and 9 as shown in . Figure 8 An embodiment of a circuit 810 (representing a load 706) using a parallel tuning circuit 805 for use with a bleeder and detune circuit 806 (i.e., an example of a bleeder and detune circuit 403 / 603) is shown. The tuning circuit 805 includes a coil 800 (e.g., a transmit coil 604) having an inductance L and a capacitor 801 having a parallel capacitance value C to form a resonant circuit having a Larmor frequency (ω 0 2 *L*C=1). Circuit 806 including switch 803 and small value resistor 804 may have the function of discharging energy and detuning coil 800. Ring-down time may be reduced when switch 803 is closed. Circuit 810 may include another switch and large value resistor arranged in parallel to 803 and 804. Therefore, for safety reasons, discharging energy may be divided into two steps. A larger resistor may be used to discharge the main energy, while a smaller resistor 804 may be used to further discharge it. When coil 800 is not used in a receive NMR / MRI signal mode, for example, circuit 810 may be adjusted to close switch 803, thereby detuning coil 800 by short-circuiting it with small resistor 804. For example, the small resistor may have an impedance value of less than 1 ohm. Circuit components 803 and 804 may be located away from coil 800 via connection to coaxial cable 802. Short cable 802 may have very small impedance and capacitor, and its effect may be negligible in low-field MRI.

[0052] Fig. 9A circuit 910 (eg, representing the load 706) including a bleeder / detuning circuit 906 and a series tuned circuit 905 is shown. A coil 900 (eg, representing the transmit coil 604) having an inductance L' and a capacitor 901 having a capacitance value C' are arranged in series to form the Larmor frequency (ω 0 2 *L'*C'=1) resonant circuit.

[0053] When Figure 7 When a voltage mode class D amplifier is used in the circuit 910, the circuit 910 may generally have a narrow bandwidth. A resistor (not depicted) may be used in series with the coil 900 to reduce the quality factor (Q) of the probe containing the coil. When RF power is sent to the coil 900, the switch 904 is closed to short the resistor 903. After sending, if the switch 904 is open, the resistor 903 may have a bleed function. In addition, except for the period of sending power, the circuit 910 may be in an open state due to the anti-parallel diode 602. In this case, the bleed / detuning circuit 906 may not be needed. In some embodiments, the circuit 910 may be used with a current mode class D power amplifier, for example, as a load 706.

[0054] In another embodiment, the load 706 ( Figure 7 ) may be an untuned coil 1004 operating as a transmitting coil 604. Fig.10 An example transmit channel 1010 using an untuned circuit with a bleeder circuit 1003 is shown. Fig.10 As shown, the frequency ω 0 The RF pulse 1000 under the condition of is amplified by the amplifier 1001, passes through the diode 1002 and is transmitted to the untuned coil 1004 operating as a transmitting coil. Fig.11 As shown in , in the case of an untuned coil 1004 as a transmitting coil, it is expected that no capacitor is used. The advantages of using an untuned coil 1004 are simplicity and flexibility. This arrangement can allow ultra-wideband and multi-frequency operation in low-field MRI. In such applications, a class D power amplifier can be arranged in close proximity to an untuned coil 1004 (e.g., a transmitting coil 604). A discharge circuit 1003 includes a switch 1103 and a low impedance resistor 1104. The switch 1103 can be closed to allow the resistor 1104 to discharge the energy on the coil 1004. The untuned coil 1004 may be blocked by an anti-parallel diode 1002, and it may not be necessary to detune the circuit.

[0055] In an embodiment, the receiving coil 203 (see Figure 2A) can be used to receive NMR / MRI signals using parallel imaging techniques. Parallel imaging techniques may reduce imaging time by several times. In order to function effectively, it may be important to minimize the interaction between the array elements of the receiving coil 203 by using decoupling techniques. In addition to geometric decoupling (e.g., as shown in FIG. 3 ), there are several decoupling techniques that can be employed, such as inductive decoupling, capacitive decoupling, and preamplifier decoupling.

[0056] Fig.12 Schematic diagram of a receiving channel using a preamplifier decoupling method. The receiving channel 1220 includes a parallel tuning probe 1200 having a parallel tuning coil operating as a receiving coil, a detuning circuit 1201, and a signal acquisition circuit 1202. The parallel tuning probe 1200 having a parallel tuning coil may have a capacitor built in. The parallel tuning probe 1200 may have a greater signal gain than a matching and tuning probe, thereby producing a lower noise figure (NF) and a higher SNR. The detuning circuit 1201 here can be used to disable the probe 1200 and protect the signal acquisition circuit 1202 when not in the receiving mode.

[0057] In receive mode, the NMR / MRI signal is amplified by a signal amplifier 1203, which may be a low noise preamplifier. The signal amplifier 1203 may include two stages (first and second stages). After the signal passes through an analog filter 1204, the signal may be amplified by a signal amplifier 1205 (third and fourth stages). This is followed by other NMR / MRI components such as a mixer 1207, an analog-to-digital converter (ADC) 1208, and a digital filter 1209. The signal is compared to a reference frequency 1206 (Larmor frequency ω 0 ) to remove the carried frequencies. In low field NMR / MRI, when ADC 1208 includes the capability to satisfy the Nyquist-Shannon sampling theorem, mixer 1207 may be replaced by a digital mixer.

[0058] Fig.13A detuning and preamplifier decoupling strategy for an MRI receive channel (e.g., receive channel 1220) is shown. A parallel tuned coil 1300 is tuned by a capacitor 1301 to be used as a probe 1200. The detuning circuit may use a switch 1305 and a small resistor 1306 (approximately equal to 0 ohms). The switch 1305 may be a MOSFET-based or PIN diode controlled by a DC bias current. An impedance transformer 1303 is used to match the impedance of the probe 1200 to the optimal noise impedance of the preamplifier 1304. In addition, the impedance transformer 1303 may also implement preamplifier decoupling. The preamplifier 1304 may have a high impedance input or a low impedance input. To implement preamplifier decoupling, the impedance of Z1 may be approximately a short circuit (from the direction of the preamplifier). In the opposite direction, the impedance of Z0 may be close to the optimal noise impedance of the preamplifier 1304. In an embodiment, several channels for collecting ambient noise may be used with Fig.13 Same setup as shown in .

[0059] Fig.14 The induced current measurement (S21) on the bench is shown, showing the first signal 1401 and the second signal 1400. The first signal 1401 is measured by connecting the circuit (impedance transformer 1303 and preamplifier 1304). The parallel tuned coil 1300 is constructed at a resonant frequency of about 5MHz. The second signal 1400 is measured without connecting the impedance transformer 1303 and the preamplifier 1304. The preamplifier 1340 is an operational amplifier with a high impedance input. The S21 measurement is proportional to the induced current on the parallel tuned coil 1300. It can be measured via a pair of decoupling sniffer loops close to the coil 1300. The difference (ΔS21) between the first signal 1401 and the second signal 1400 indicates that the preamplifier decoupling strategy can reduce the coil current by approximately 19dB.

[0060] It will be apparent to those of ordinary skill in the relevant art that other suitable modifications and adjustments may be made to the methods and applications described herein without departing from the scope of any embodiment. It should be understood that, although certain embodiments have been shown and described herein, the claims are not limited to the specific form or arrangement of the parts described and shown. Illustrative embodiments have been disclosed in this specification, and although specific terms are used, they are only used in a general and descriptive sense, not for the purpose of limitation. In view of the above teachings, modifications and changes to the embodiments are possible. Therefore, it should be understood that the embodiments may be practiced in a manner different from that specifically described. All disclosures, patents, and patent applications mentioned in this specification are incorporated herein by reference to the extent that each individual disclosure, patent, or patent application is specifically and individually indicated to be incorporated by reference.

Claims

1. A method for stabilizing the magnetic field of a mobile MRI device using a permanent magnet, the method comprising: providing a main static magnetic field by the permanent magnet of the mobile MRI device; A frequency measurement module of the mobile MRI device provides a radio frequency pulse to induce a free induction decay (FID) signal; The frequency measurement module monitors the phase of the FID signal to determine the frequency of the main static magnetic field; The frequency of the main static magnetic field is adjusted by a compensation module. 2 . The method of claim 1 , wherein the frequency measurement module comprises at least one channel for field monitoring.

3. The method of claim 1, wherein the frequency measurement module comprises an NMR probe to monitor the field.

4. The method according to claim 1, wherein the frequency measurement module comprises 13 C or 19 F as NMR sample. The method of claim 1 , wherein the frequency measurement module comprises a monitoring coil.

6. The method of claim 1, wherein the compensation module comprises at least one channel having a shim amplifier and a shim coil, and Wherein adjusting the frequency of the main static magnetic field includes adjusting a current applied to the shim coil.

7. A system for transmitting an RF field in a mobile MRI apparatus, the system comprising: The sending channel includes: Class D power amplifier, Anti-parallel diode, One or more circuits configured for detuning or bleed, and At least one of the RF coil is tuned and the RF coil is untuned.

8. The system according to claim 7, wherein The class D power amplifier comprises a voltage mode class D power amplifier, and the system further comprises: Parallel tuned probe; at least one detuned circuit; as well as At least one bleeder circuit.

9. The system according to claim 7, wherein at least one of the tuned RF coil and the untuned RF coil is a series tuned RF coil; and and The transmit channel further includes a resistor connected in series with the series tuned RF coil to increase bandwidth.

10. The system of claim 9, wherein the class-D power amplifier is a current-mode class-D power amplifier.

11. The system of claim 7, wherein the class-D power amplifier is a voltage-mode class-D power amplifier and at least one of the tuned RF coil and the untuned RF coil is an untuned RF coil, and the system further comprises a bleeder circuit.

12. A mobile MRI device, comprising: a permanent magnet configured to provide a main static magnetic field; as well as A receiving channel, comprising: Parallel tuned probe, Preamplifier; at least one detuned circuit; an impedance transformer; and Signal acquisition device.

13. The mobile MRI device of claim 12, wherein the preamplifier comprises a high impedance input.

14. The mobile MRI apparatus of claim 12, wherein the preamplifier comprises a low impedance input. 15 . The mobile MRI apparatus according to claim 12 , wherein the receiving channel is configured to collect noise for noise cancellation.

16. The mobile MRI apparatus of claim 12, further comprising one or more detuning circuits.

Images