Arbitrary core distributed active programmable transmit MRI coil

Abstract

A magnetic resonance imaging transmitting coil comprises: an input for receiving one or more control signals; one or more voltage or current sources for generating one or more source signals; at least two coil segments, each coil segment for receiving one or more control signals and one or more source signals, each coil segment having one or more switches to convert the one or more source signals to one or more arbitrary frequency currents, the one or more currents causing a magnetic field.

Description

Technical Field

[0001] This disclosure relates to magnetic resonance imaging (MRI) coils, and more specifically to MRI coils that can be used with any nuclei. background

[0002] There are 118 known elements. Almost all of these elements have NMR (nuclear magnetic resonance) active isotopes, and at least 39 different nuclei have been shown to be biologically relevant. Nevertheless, most MRI scans today are based on only one nucleus— 1 H.

[0003] 1 H has the second highest gyromagnetic ratio, only in 3 Following H, and being the most abundant in the human body, it has historically been the preferred imaging tool when SNR (signal-to-noise ratio) is insufficient. However, over the decades, MRI technology has made significant progress in SNR through hyperpolarization, high-field and ultra-high-field magnets, anatomically conformal receiver coils, improved reconstruction, and other techniques. Utilizing these SNR gains, in addition to… 1 Imaging of nuclei other than H or X nuclei has become more clinically feasible, and various studies have emerged utilizing the near-perfect nuclear specificity of NMR / MRI to obtain images using only... 1 Information that H could not obtain. But ultimately, the dissemination of this research was still hampered by the low availability of the tools that enabled it.

[0004] Nuclear specificity is a double-edged sword in terms of ease of conducting experiments. Studying nuclei X requires significant investment to acquire additional, expensive MRI hardware, such as RF amplifiers, coils, and receiver chains, to enable imaging of specific nuclei of interest. This high barrier to entry severely limits research on many nuclei that, due to their low natural abundance in the human body and thus limited general applicability, may have more direct biological relevance. Examples include, but are not limited to, those involving nuclei X. 23 Na、 31 P, 13 C and 2 H. 2 H has very low natural abundance, but has recently regained research interest as a metabolic probe upon ingestion or injection. Other examples include... 3 He or 129 Xe (which is used for functional lung imaging during hyperpolarization and inhalation) and 7 Li (which is used to treat bipolar disorder and fractures). In addition to these and other unlisted elements, there are 39 different nuclei from 33 elements with untapped potential, which are biologically related to these elements.

[0005] Coils for the X-core are not widely available. Even if one were obtained, further difficulties arise. If using... 1 To obtain structural images and / or co-register multiple images from different nuclei using H, changes to the coil introduce complexity and registration issues into the MRI workflow.

[0006] Efficiently exciting arbitrary nuclei at high fields using a single coil has remained an unsolved problem in macroscale MRI. The usual approach to multinucleus studies with a single coil involves multituning the coil. However, complexity increases very rapidly with the number of nuclei, and quality decreases with more tuning due to increased losses. This leads to a general limitation of multituned coils to only two or three nuclei. Recently, an impressive five-fold tuned coil has been described for... 1 H, 19 F, 31 P, 23 Na and 13 Imaging with C is possible, but adding more nuclei would be challenging. Some other multi-nucleus coils attempt to use variable circuitry to tune the coil to different nuclei on the fly, but they all have limited tuning ranges and may require mechanical modifications for better performance. NMR has a small length scale, and some reported methods use untuned microhelices and emitter wire coils with resistive terminations, but these cannot be extended to macroscale MRI. Brief description of the attached diagram

[0007] Figure 1 shows a comparison between the current method of connecting the coil and the method of connecting the coil according to an embodiment.

[0008] Figure 2 A schematic diagram of a simulation for an arbitrary core coil is shown.

[0009] Figure 3 A schematic diagram of an embodiment of an arbitrary core coil is shown.

[0010] Figures 4A-4F Simulation curves of current versus switching frequency at base frequency are shown for different coil inductance, resistance, and control signal timing.

[0011] Figures 5A-5C The simulation results for different input voltages are shown in graphs.

[0012] Figures 6A-6D Photographs and graphs of the results for an embodiment of an arbitrary core coil are shown.

[0013] Figure 7 This demonstrates the use of arbitrary core coils for 1 H, 23 Na、 2H and 13 MR embodiment of C.

[0014] Figure 8 An image of an ex vivo MRI using an embodiment employing an arbitrary nuclear coil is shown.

[0015] Figures 9A-9B It shows 1 H arbitrary core coil B1 + The image of the figure.

[0016] Figures 10A-10D Examples of saddle-shaped and birdcage-shaped coils are shown.

[0017] Figure 11 An example of a control waveform used to generate current in a coil is shown.

[0018] Figure 12 The control and current waveforms in the coil, as well as the current frequency, are shown. Detailed Implementation

[0019] This embodiment addresses the issues of X-nucleus coil availability and complexity by focusing on an MRI system capable of imaging arbitrary nuclei using a single set of hardware. The embodiment concentrates on RF emission and proposes a novel MRI emission coil capable of exciting resonances of arbitrary nuclei on a macroscopic scale within a whole-body MRI scanner.

[0020] The embodiments presented here depart entirely from the concept of tuned macroscopic-scale coils and propose an untuned but scalable 9 cm diameter transmit surface coil capable of efficiently transmitting at arbitrary frequencies using a unique architecture comprised of numerous high-frequency semiconductor power switches directly integrated into the coil structure. In doing so, the coil and RF amplifier are combined into a single device that directly converts DC power into an RF field at any relevant frequency. On the receiver side, the capabilities of the transmit coil can be demonstrated using several single-tuned receiver coils.

[0021] Inductance resists changes in current that cause changes in magnetic field. Inductance makes it difficult to generate rapidly oscillating RF magnetic fields. As an example, in the context of human-scale MRI, a single-turn toroidal coil with a diameter of 9 cm can have an inductance of approximately 300 nH (nahens). According to the Biot-Savart law, approximately 360 mA of current is required to generate a moderate 5 μT magnetic field at the center of this toroidal coil. The Biot-Savart law calculates the current at the location... r Magnetic field generated by electric current B .

[0022] The formula used for inductors is ,in v It's voltage. L It is an inductor, and It is the derivative of current with respect to time. In order to... 1 The H 3T Larmor coil, operating at 127.7 MHz, generates a peak current of 360 mA in a 300 nH coil, requiring a peak voltage of approximately 87 V at 127.7 MHz. To achieve this, for... 1 Similar excitations of nuclei other than H typically occur at lower frequencies, but due to the lower gyromagnetic ratio, the required magnetic field and current increase proportionally. The required voltage for... 1 The H and X nuclei are approximately identical to produce equal flip angles over equal durations. To achieve these large voltages, a resonant circuit is typically used to convert the lower voltage from the output of the RF power amplifier into the voltage required to generate current in the inductor. However, unfortunately, such a resonant circuit has an inherently limited bandwidth, constrained by the Bode-Fano limit. In a typical MRI context, these bandwidths are on the order of hundreds of kHz, capable of covering the resonance of only one nucleus at a time.

[0023] Because resonant circuits are severely limited in their bandwidth, to achieve multi-core capability, one might imagine simply connecting a very high-voltage amplifier directly to an untuned coil to generate around 87 volts. However, this becomes difficult at RF frequencies, as RF amplifiers capable of reaching 87 volts are typically quite large. Furthermore, if the transmitter line separates the RF amplifier from the untuned coil, this would cause significant power reflections from the coil, potentially damaging the amplifier and leading to safety issues.

[0024] Alternatively, instead of using a conventional 50-ohm output RF amplifier, a transistor can be used to switch the voltage directly at the coil. In fact, this is essentially the approach taken by Mandal et al. and Hopper et al. in their ultra-broadband low-frequency NMR system covering up to 3 MHz. (Mandal S, Utsuzawa S, Cory DG, Hürlimann M, Poitzsch M, Song YQ. An ultra-broadband low-frequency magnetic resonance system.) Journal of Magnetic Resonance .2014;242:113-125. doi:10.1016 / j.jmr.2014.02.019; Low-frequency NMR with a non-resonant circuit by Hopper T, Mandal S, Cory D, Hürlimann M, Song YQ. Journal of Magnetic Resonance.2011;210(1):69-74. doi:10.1016 / j.jmr.2011.02.014; and Mandal S, Utsuzawa S, Song YQ’s An extremely broadband low-frequency MR system. Microporous and Mesoporous Materials .2013; 178:535.oi:10.1016 / j.micromeso.2013.03.040.) In this method, the H-bridge alternately switches from one side of the coil to the other at an RF frequency of up to 200 volts, thereby generating an RF magnetic field. Figure 1A A circuit architecture with coil 10 and DC voltage source 20 is shown. The voltage source is connected to the coil via a set of switches 12, 14, 16, and 18. This represents a conventional H-bridge. The switches connect the DC voltage source to the coil in alternating directions depending on which set of switches is closed. Figure 1B A first configuration is shown, in which switches 12 and 18 are closed to connect a DC source in one direction. When switches 12 and 18 are open and switches 14 and 16 are closed, an AC current is generated in the coil. The current flowing in the coil generates a magnetic field.

[0025] Ultimately, however, this approach is limited by the capabilities of the semiconductor devices used. Devices capable of handling higher voltages tend to be slower, and as of this writing, the inventors are unaware of any device capable of handling frequencies up to 127.7 MHz in a reasonable manner. 1 At the H 3T Larmor frequency, the coil interface is directly connected without a resonant circuit. As Mandal et al. noted regarding their method, "this technique can be extended to higher frequencies if smaller coils are used." While this is effective and similar to the approach employed with ultrawideband microhelical coils for multinuclear NMR, too many coils are required to extend to high-field MRI at the human scale.

[0026] exist Figure 1D In the embodiment shown, the simplest configuration of the Arbitrary Core Distributed Active Programmable Emitter (ADAPT) coil divides the coil into two halves, half 22 and half 24. Two switches 26 and 28 alternately connect a DC voltage source 20 to each half. Figure 1E The diagram shows that when switch 26 is closed, current flows through coil half 24. Figure 1F The diagram shows that when switch 28 is closed, current flows through coil half 22. Changing the connection in the coil generates AC current, similar to an H-bridge.

[0027] The advantage of the ADAPT coil configuration is that it allows for scalable coil splitting by adding more switches and connections to the DC voltage source, such as... Figure 1G-Figure 1I As shown. By dividing the coil into smaller segments, the same voltage from the voltage source can drive more current through the inductance of the coil segment without adding more voltage stress to the switch. Figure 1G An embodiment with three coil segment pairs 30, 32, and 34 is shown. Each segment pair has a pair of switches. Figure 1H In the diagram, one switch in each pair is closed, switches 36, 38, and 40, and current flows in the direction of the current arrow. Figure 1I The opposite switches, 42, 44, and 46, are shown closed in each pair. This allows current to flow in relation to... Figure 1H It flows in the opposite direction to the flow direction.

[0028] Figure 1G-Figure 1I The embodiments provided a solution for the limited voltage of semiconductor devices. Instead of splitting the coil in half, the coil can be broken down into smaller segments, each with a set of switches connected to the negative terminal of a DC voltage source. Now, the same voltage source drives several smaller inductors, coil segments connected in parallel, and can generate a larger RF current. These RF currents then together produce a larger magnetic field than the magnetic field produced by a set of switches individually. Segmenting the toroidal coil scales approximately linearly with the radius of the toroidal coil, or proportionally with its circumference. Conversely, using many parallel micro-spiral coils to form an effectively larger toroidal coil will be proportional to the area of ​​the toroidal coil, or proportional to the square of its radius.

[0029] The embodiments described herein enable the direct application of voltage to the coil, unrestricted by the semiconductor voltage capability. The embodiments utilize a method similar to that of a distributed active transformer, which addresses a similar problem of generating higher output voltages from low-voltage CMOS transistors by placing several transistors in a loop. Figures 1D-1F It shows the most basic case and allows for... Figures 1A-1C The H-bridge in the comparison is used.

[0030] Similar to the H-bridge, alternating switches at RF frequency change the current path of the DC voltage source. Instead of using four switches as in the H-bridge, only two switches are used, at the cost of driving only half of the coil at a time. Importantly, however, the two switches not connected to the negative terminal of the DC voltage source are no longer used. This eliminates the complexity of driving the "high-side" power switch, whose reference voltage would shift with a large voltage across the coil. Using only the "low-side" switch makes faster switching available. It should be noted that this embodiment uses one switch per coil segment, but more switches can be used to achieve various trade-offs. For example, parallel switches can increase current handling capacity at the expense of increased parasitic capacitance, or series switches can increase voltage handling capacity at the expense of increased on-resistance. "High-side" switches can also be added if a sufficiently fast switch becomes available.

[0031] Driving only half of the coil at a time results in some DC current in each coil segment. This can cause some static magnetic field inhomogeneity during RF transmission. Increasing the number of coil segments reduces static magnetic field inhomogeneity because adjacent coil segments with opposite DC currents become smaller and closer together. This also applies to DC current flowing from a DC voltage source through the center of the coil. However, if static magnetic field uniformity during RF transmission is more critical, adding a second identical coil on top of the first coil can potentially eliminate these effects. Rotating the second coil, which is one coil segment in length, causes the RF current to flow in both directions throughout the loop. The magnetic field from the DC current will then be in opposite directions and canceled out.

[0032] Another issue to address lies in the characteristics of the switch, particularly the on-resistance and parasitic parallel capacitance. These non-ideals degrade the performance of the ADAPT coil by increasing coil losses. On-resistance increases losses as current flows, and capacitance increases losses as charge on the capacitor is transferred to ground when the switch is closed. These loss mechanisms must be considered when designing the ADAPT coil.

[0033] Directly switching the DC voltage across an inductor does not produce a perfect sine wave. This introduces harmonics at the fundamental switching frequency, and the amplitude depends on the exact switching mode and the circuit inductance and capacitance, especially parasitic effects. These harmonics have no effect on the MRI signal and should be much smaller than the fundamental frequency. However, they can increase SAR (specific absorption rate), and their safety should be considered. Several harmonic cancellation techniques used in power electronic inverters can remove harmonics without resonant circuits or traditional frequency filters. These techniques can be implemented by stacking more ADAPT coils on top of each other, or by using more complex control signals.

[0034] In the simulation, one embodiment uses a gallium nitride (GaN) power switch, chosen for its very low on-resistance (0.04 ohms) and fast switching speed exceeding 100 MHz, for practical simulations involving switching defects. The simulation includes an additional capacitor and series resistor placed in parallel with the GaN power switch to represent a transient voltage suppression diode to protect the power switch from damaging overvoltages. Figure 1D The equivalent circuit is shown, where the inductance between the DC voltage source and the point where the coil segment is divided is set to 100nH, which is referred to here as the source inductance.

[0035] Figure 2 A circuit diagram fabricated in an Advanced Design System (ADS) is shown. SRC1 52 represents a DC voltage source providing power to the resulting RF magnetic field, SRC2 54 represents a power supply for a power switching integrated circuit or chip, X1 56 and X2 58 represent power switches, SRC3 60 and SRC4 62 represent digital control signals for the switches, L1 64 represents the inductance from the DC source to the coil segment (source inductance), L2 66 and L3 68 represent two complementary coil segments, and R1 70 with C2 72 and R2 74 with C1 76 represent parasitic effects from transient voltage suppression diodes. It should be noted that while this embodiment uses digital control signals, other embodiments may use analog control signals. In the case of analog control signals, the switch can also act as a more continuously varying variable resistor. It should be noted that in the case of analog control signals, the switch behaves like a standard amplifier and can be implemented as such. The term "switch" as used herein includes amplifiers, diodes, relays, transistors, and variable resistors. The probes for the current are positioned at I_probe 1 78 and I_probe 2 79.

[0036] The inductance of the coil segment was scanned for a constant DC voltage source to show the effect of increasing current as the coil segment inductance decreases. To illustrate the effect of the source inductance, the source inductance was also scanned for both the constant segment inductance and the DC voltage source. The source and segment resistances were also scanned, and the skew between the two complementary control signals and their rise / fall times were also scanned. While the given example uses two control signals, other embodiments could use only one control signal from which other control signals are derived, for example, by connecting switches of opposite polarity. Finally, the DC voltage was also scanned for the constant coil segment and the source inductance for several frequencies of interest. Due to the symmetry of the ADAPT coil, only one pair of coil segments needs to be simulated, and the reduced simulated segment inductance represents dividing the coil into more segments. Harmonic balance simulation in ADS was used for a faster simulation speed than transient simulation, and the simulation data from ADS was imported into MATLAB (The MathWorks, Natick, USA) for plotting.

[0037] Figure 4 shows the results of ADS simulations of the ADAPT coil operating between 10 MHz and 130 MHz for several different coil segments and source inductance values. As the inductance of each coil segment decreases, as... Figure 4A As shown, the increased current flowing through each coil results in a higher magnetic field, demonstrating one of the benefits of dividing the coils into more segments in the ADAPT structure. Interestingly, a drop is observed across various frequency ranges. This occurs when the circuitry and switching parameters happen to convert more DC voltage into a DC field or frequency harmonics rather than the desired fundamental frequency. Figure 4B The current for various source inductors is shown. Figure 4C The current for various resistor segments is shown. Figure 4D The current is shown for various source resistors. Figure 4E The current for various control skews is shown, and Figure 4F The diagram shows the current for various control rise / fall times. These illustrate the effects of non-ideals in the circuit and non-ideal control signals.

[0038] However, as Figures 5A-5C As shown, the decrease in the curve in Figure 4 does not necessarily indicate a lower magnetic field strength, because as the DC source voltage increases, the ratio between the fundamental current and the harmonic current becomes better and can overcome the decrease. This effect occurs at higher voltages because when the drain voltage on the switch is sufficiently negative, reverse conduction through the switch occurs and parasitic oscillations that contribute to harmonics are blocked. When the source inductance changes, as... Figure 4B As shown, the effect on the current is much smaller because the source inductance primarily carries DC current. This assumption becomes even better with higher source inductance, but changing the source inductance mainly alters the position of the descent in the curve.

[0039] Figures 5A-5C Results of ADS simulations for several cores of interest in the time and frequency domains are shown, where the DC source voltage is scanned to vary the RF amplitude, and the segment inductance and source inductance are kept constant at 10 nH and 100 nH, respectively. The results for the nuclei of interest in the time and frequency domains are also shown. 1 H, 13 C and 15 A simulation diagram of the ADAPT coil operating at a 3T Larmor frequency (N) was chosen as an example with good frequency spacing. Note the cases with a 10 nH segment inductor and a 100 nH source inductor. 13 The C frequency falls on a decrease in the curve shown in Figure 4.

[0040] As the operating frequency decreases, the time-domain current waveform and the resulting magnetic field become less sinusoidal. This is mirrored in the frequency domain plot. Higher frequencies are closer to a sinusoidal waveform due to filtering by inductors and capacitors. The switch drain voltage limits the maximum RF output because if the voltage exceeds a threshold, such as 40 V for the GaN switch used here, the switch may open. As can be seen from the plot, lower operating frequencies also tend to reach this limit at lower DC source voltages. Ultimately, however, in this embodiment, the maximum achievable current and RF magnetic field are similar across different frequencies. The efficiency, calculated as the current at the fundamental frequency divided by the square root of the power from the DC voltage source, is also similar across different frequencies at higher field strengths. This current efficiency is specific to a particular coil segment with these circuit parameters, but the magnetic field efficiency can be found by assuming the coil geometry and knowing how many coil segments are needed to create it. If eight 10 nH coil segments can produce a coil with a 9 cm diameter, then... Figures 5A-5C The efficiency shown can be achieved by... Value and Approximation Multiply at the center of a 9 cm diameter coil from Convert to Of these, 14 comes from Biot-Savart's law, used to convert the current in a 9 cm toroidal coil into a magnetic field at its center, and Eight times the power from eight coil segment pairs, but expressed as current efficiency in square root form. Efficiency is expected to improve further when using more complex control signals, especially at lower RF frequencies.

[0041] The reduced segment inductance also reduces the switching voltage stress for any given current, thereby reducing the magnetic field, because... .

[0042] The inventors used commercially available components assembled on a commercially manufactured four-layer FR4 printed circuit board (PCB) from JLCPCB (Shenzhen, China) as shown in Figure 6 to implement the simulation architecture. Figure 3 A block diagram showing the chip used and its connections, as well as auxiliary devices for operating the ADAPT coil, is illustrated. The diagram shows a pair of segments on the coil, which will replicate any number N for the paired coil segments 108. The coil can be further replicated.

[0043] The MRI scanner 82 and waveform generator 80 provide inputs to two low-voltage differential signal (LVDS) digital converters 84 and 86, which independently control two sets of out-of-phase high-power switches 100 and 102. Having two independent inputs allows for the simultaneous shutdown of both sets of switches when neither is transmitting nor generating current. An LVDS repeater 88 then replicates these two inputs to each associated switch. This embodiment uses GaN switches. Related chips for digital logic conversion, such as an LVDS-to-CMOS converter 96, are also added, along with voltage and power regulation using regulators 90, 92, 94, and all associated decoupling capacitors. Transient voltage suppression (TVS) diodes 104 and 106 are placed across each switch to protect them in the event of exceptionally strong transients.

[0044] To operate the coil, a pulse sequence from the MRI scanner 82 triggers an externally generated LVDS signal that sets the desired RF frequency. The RF amplitude is set by the output voltage 92 of the DC-coupled audio amplifier, which acts as the main DC voltage source for converting the RF. Only hard RF pulses are used in this demonstration; however, it should be noted that the DC voltage and RF frequency can be time-varying. It should also be noted that the source here includes a voltage source, but it can also be a current source. Whether it is a voltage or current source, the output of the source will be referred to as the source signal. It should be noted that although a single source is used here, multiple source signals can be used instead of a single source signal and routed to different coil segments.

[0045] In other embodiments, RF amplitude modulation can be achieved by making the output of the audio amplifier time-varying. It should be noted that the amplifier here is an audio amplifier, but amplifiers or sources with frequencies outside the audio frequency range can also be used. The amplifier bandwidth may affect the transient characteristics of the coil.

[0046] Because the coil is untuned and its self-resonant frequency is higher than any operating frequency, by designing the coil segment with sufficiently low inductance to achieve high resonance with the switch parasitic capacitance, the coupling between the transmitting and receiving coils is negligible, and no additional transmitting coil detuning is required during reception.

[0047] A benchtop DC voltage scan was performed at six operating frequencies to empirically characterize the relationship between the DC supply voltage (400 mV to 6000 mV in 400 mV increments) from the DC-coupled audio amplifier and the amplitude of the RF magnetic field output by the fabricated ADAPT coil. The frequencies were chosen as 12.94, 19.60, 32.11, 33.78, 51.69, and 127.7 MHz, corresponding to 3T... 15 N、 2 H,13 C 23 Na、 31 P and 1 H-core. To measure the amplitude of the RF magnetic field, a calibrated electromagnetic compatibility probe (Beehive Electronics, Sebastián, USA, Model 100B) located 19.6 mm from the center of the coil was used in conjunction with the real-time spectrum analyzer mode of a FieldFox N9918A (Keysight Technologies, Santa Rosa, USA). A 25 ms RF pulse was sent to the input of the sine wave to LVDS converter every 500 ms using an SDG6022X function generator (Siglent Technologies, Shenzhen, China).

[0048] Figures 6A-6C An image of the coil manufactured in this embodiment is shown, and Figure 6D The graph shows the results of the benchtop voltage scan. Figure 6A A top view of coil 109 is shown, in which... Figure 6B A close-up view of connector 110 is shown. Figure 6C The back of coil 109 is shown. The coil elements follow... Figure 3 The block diagram shows a layout with eight coil segment pairs. Figure 1G The version mentioned above. Figure 3 A schematic diagram of a single coil segment pair is shown. Although this embodiment uses eight, it is not intended to, and should not, imply a limitation on that number.

[0049] Figure 6D The results of a benchtop voltage scan are presented, demonstrating the relationship between the probed RF output and the DC voltage from the audio amplifier. A positive and approximately linear correlation between the amplitude of the RF output and the DC voltage is evident across the scanned voltage range. While consistent with the trend depicted in Figure 4, the experimental results reflect a subtle difference not captured by the simulations discussed above. It is noteworthy that... Figure 6D It can be observed that the empirical amplitude of the RF field begins to saturate at higher DC voltages along the scan range, particularly for 51.69 MHz and 127.7 MHz.

[0050] exist Figure 7 At the top, from left to right, are vials from the sample collection. 1 H, 2 H, 23 Na and 13 Image C. The sample consisted of distilled water, saline solutions with three different salt concentrations, deuterium oxide, and urea. 13 C) Composition. Figure 7 The lower left corner shows a combined image of the four cores. Figure 7 The lower center shows an optical image of the sample. Figure 7 A diagram of the sample is shown in the lower right corner.

[0051] Figure 8 Showing for 1 H and 23 Ex vivo images of bone-in ham ribs and pig knees from Na. This demonstrates the imaging results of the actual tissues. The experiments below provide more information about these images.

[0052] experiment Without replacing the ADAPT coil, the phantom was photographed by changing the control signal used for transmission. 1 H, 2 H, 13 C and 23 Na image. The phantom consists of five 10 mL bottles arranged in a circle and a centrally located... 13 A 1 mL tube containing urea (8 M concentration) and gadolinium (0.1 mM concentration) enriched in C was used. Figure 7 The center bottom is shown. Five 10 mL bottles consist of distilled water, deuterium oxide (Sigma-Aldrich, St. Louis, USA), and saline solutions with three different NaCl concentrations (10 g / L, 20 g / L, and 40 g / L). Because the scanner used does not have multi-core capability, a custom receiver chain was built for a single-channel receive-only coil. This receiver chain upconverts the X-core signal to... 1 H frequency to use existing 1 The H-receiver chain is used for data acquisition. Individual single-tuned surface receiver coils are used for imaging each nucleus.

[0053] Pulse sequences were written using a Spin bench / RTHawk platform (Vista.ai, Los Artos, USA) and run on a GE 3T MR750w scanner (GE Healthcare, Waukeshaw, USA). Gradient echo (GRE) sequences were used with the following parameters: coronal imaging, 3.2 ms non-selective hard RF pulses (except for those with 2 mm slice selection). 1 H), 300ms TR, 3.7ms TE, 6 V DC voltage emitter coil input for all cores (except those with 200 mV). 1 H); for 1 H is 1.25×1.25mm 2 Resolution, 256 × 256 matrix, and 125 kHz readout bandwidth; for 2 H 10×10 mm 2Resolution, 16×16 matrix, 25 kHz readout bandwidth, and 100 averages; for 13 C and 23 Na, 5×5 mm 2 Resolution, 32×32 matrix, and 50 kHz readout bandwidth, with averages of 100 and 25, respectively. Based on benchtop B1 calibration using the same equipment and cables as the imaging experiments, the flip angle at the center of the phantom was determined for... 1 H, 2 H, 13 C and 23 Na was estimated at 3.5°, 10°, 53°, and 63°.

[0054] They also filmed bone-in ham steaks and slices of pork knee outside the body. 1 H and 23 Na images, such as Figure 8 As shown. The parameter deviations from phantom imaging are: 1) non-selective 1 H-stimulus, and 2) for 23 Na imaging, 100 and 500 average values ​​of images of ham ribs and pig knees, respectively, have a TR of 100 ms.

[0055] To show, 1 The H image was cropped to 128×128, and the X kernel image was interpolated to 128×128 using zero-padding and a Hamming window.

[0056] In order to map 1 H B1 + The experiment used a 3-mm thin-film phantom (30 mL of 1 g / L CuSO4 and 50 g / L NaCl solution, approximately 260 ms T1), placed 27 mm from the ADAPT coil. The sequence parameters were: 1 s TR, 3.7 ms TE, 2.5 × 2.5 mm. 2 Resolution, 128×128 matrix, non-selective hard pulse excitation, triple averaging. All images are received using a body coil, while excitation is achieved using either the body coil or the ADAPT coil without moving any hardware. When the body coil is used for transmission, the ADAPT coil is in place and energized (but does not transmit). Two images excited using the body coil. and by and 2 Shoot from a flip angle. C(r) The magnetization is received by the body coil with sensitivity weighting, and complete relaxation with magnetization having a long TR is assumed. [1] [2] [3] as well as [4] Then, given any image It uses the same setup for data acquisition, but uses an ADAPT coil for excitation. [5] and [6] Formula [6] is used to calculate B1 for the ADAPT coil for a series of voltages (up to 10V) shown in Figure 9. + Figure. Due to the nonlinearity of the ADAPT coil, the DC voltage cannot be doubled to double the flip angle. No reference emitter coil was used for other nuclei in this experiment.

[0057] Figure 9A The 1H B1 is shown for different voltages. + The figure provides insight into the causes of RF field saturation discussed above. In the current coil implementation, the higher B1 field appears to disrupt the phase relationship between switches at different locations within the coil, resulting in some form of destructive interference. This interference is likely caused by the B1 field and feedback from the control lines to the switches, influenced by asymmetrical routing of the control lines and Ethernet connectors. Future coil implementations will place greater emphasis on coil layout symmetry and shielding of the control lines. Figure 9B It shows that for Figure 9B The flip angle of the four regions of interest shown on the left is a one-dimensional plot of the DC voltage.

[0058] In this way, arbitrary nuclear emitter coils can be constructed for macroscale MRI. To cover more anatomical structures, emitter coil arrays can be fabricated directly using the surface coils presented here. Because coil tuning and matching are not involved, array coils using ADAPT coils are expected to be easier to construct than in the usual case where coil coupling affects their tuning. Because resonance is not involved, the effect of patient load relative to coil inductance should also be negligible.

[0059] The same concept of scalability can be applied to coils of other shapes instead of surface coil arrays, including solenoids and gradient coils. Birdcage coils can be constructed by dividing each RF current-carrying segment of the birdcage into smaller inductors, just as was done for the surface coils presented here. Figures 10A-10D Different configurations for "saddle-shaped" and "birdcage-shaped" coils are shown. Figure 10A A single saddle-shaped coil 112 is shown, each segment having a connector such as 118, which can be connected to a source... Figure 6B Connectors 110 are the same as or of different types. Figure 10C A saddle-shaped coil 116 with more vertical segments is shown, and Figure 10B A birdcage-shaped coil 114 with a horizontal segment forming a circle is shown. Figure 10D An embodiment of a coil with connecting wires is shown. It should be noted that although a specific shape of coil is mentioned in this discussion, coils of any shape can be used.

[0060] Higher output current and frequency should be possible by making each coil segment have even smaller inductance. Higher frequencies can also be achieved by selecting switches with smaller parasitic capacitance, but this usually comes at the cost of higher on-resistance.

[0061] Amplitude and frequency modulated pulses can be achieved by changing the DC voltage source to a time-varying voltage source for amplitude modulation and changing the switching frequency for frequency modulation over time. Due to the nonlinearity of the coil, careful characterization and pre-compensation of the coil may be necessary to improve waveform accuracy. Figure 11 Using data from Figure 5 and from Figure 2 An example of this situation is shown in the simulation of the circuit model.

[0062] In a given Figure 2 Given the circuit model, the input voltage waveform and control pulses are optimized to generate a Hamming window synchronization pulse with a time-bandwidth product of 4. The top row represents the ideal pulse, and the last column represents the ideal frequency distribution. Figure 11 The data displayed from left to right includes the input voltage, control phase, output current, and output frequency, as well as a magnified view of the output frequency for each pulse. The second row shows the input of an ideal pulse. Figure 2 The result of the circuitry in the diagram leads to a severely distorted frequency distribution. The third line shows the correction for the input voltage only, and the fourth line shows the correction for both the input voltage and the control phase. Control phase correction is necessary because the phase of the output current changes slightly with variations in the input voltage due to the nonlinearity of the actual switching model. RF amplifiers can shape pulses based on the nonlinearity of the coil.

[0063] Figure 12 This shows the direct transmission to Figure 2 The switch in the circuit model is at 32.1MHz. 13 C and at 127.7 MHz 1 The average control waveform of H. Figure 12 The second row also shows the current waveform, and the third row shows the current, which is related to the magnetic field. The simulation also generated many intermodulation frequencies. The DC voltage used for the simulation was set to 0.5V. This example demonstrates that current can be generated simultaneously at multiple desired frequencies.

[0064] Benchtop testing shows that the ADAPT coil can also be used as an arbitrary core receiver, but its large SNR penalty in its current implementation makes it suboptimal. Mainly, the on-resistance of the switch, harmonic conversion, and switching losses contribute more to the coil noise than is typically the case with copper and passive components alone, especially when the signal on the coil is low, as shown by the low efficiency at low current in Figure 5. Other potential paths to a broadband receiver coil include using a resonant coil with a broadband high-impedance amplifier or an amplifier with low noise, eliminating the need for a resonant amplification stage, possibly with a cryogenic amplifier.

[0065] Even without pairing with an optimal arbitrary core receiving coil, the ADAPT coil itself offers numerous benefits. For the standard... 1 H research suggests that ADAPT coils could potentially significantly reduce the production cost of MRI emitter coils by eliminating the need for any specialized manual assembly. This could be particularly attractive for emerging low-cost MRI scanner initiatives. All components are commercially available, and coils can potentially be ordered fully assembled from standard PCB facilities. When parts are purchased in bulk, unit costs decrease, further reducing the cost of the MRI system. Furthermore, using an audio amplifier instead of a more expensive and often less power-efficient RF amplifier to provide RF power reduces costs.

[0066] Unlike standard MRI coils that can only operate for a single nucleus at a single static field strength, the ADAPT coil can operate for the same nucleus or any other nucleus at any static field strength. It can be reused between MRI scanners with different static field strengths, and can also be used at different static field strengths in MRI scanners with variable static field strengths (which vary spatially or temporally).

[0067] In addition to lower cost, the ADAPT coil also offers easier transmit B1 calibration due to negligible load effects, and because the current in the coil stops almost immediately when the switch is turned off, it also provides a much lower dead time between transmit and receive times in MRI systems used for ZTE / UTE studies. In one embodiment, this reduction in dead time occurs because the transmit coil can operate with shorter transmit pulses without resonant circuitry.

[0068] Other variations include programming or otherwise configuring control signals to cause the voltage or current source to generate a source signal in the form of current with a desired waveform and desired duration. One or more control signals can be configured to cause the voltage or current source to generate currents at multiple desired frequencies to simultaneously excite multiple cores.

[0069] The control signal can be configured to cause a voltage or current source to generate a low-frequency current in the range from DC to kHz for multiphoton excitation, spatial coding, or field shimming. The control signal can also be configured to cause the voltage or current source to simultaneously or sequentially generate a low-frequency current in the range from DC to kHz and a high-frequency current in the range of MHz for various applications. For example, low-frequency field shimming can be achieved while using a high-frequency excitation core. kHz and MHz magnetic fields generated by the current can also be used together to generate multiphoton excitation on one or more cores.

[0070] The transmitting coil can be combined with the receiving coil for use as described above. Figure 12 The various control signals discussed can simultaneously or sequentially generate X-ray images of arbitrary nuclei. The transmit coils can be used at multiple MRI scanners with different static magnetic fields for the same or different nuclei.

[0071] The NMR signal excited by the transmitting coil can also be used for NMR spectroscopy, detection of specific molecules via NMR spectroscopy, spatial localization NMR spectroscopy (MRS), and magnetic resonance imaging (MRSI).

[0072] Furthermore, this written description refers to specific features. It should be understood that the disclosure in this specification includes all possible combinations of these specific features. For example, where a specific feature is disclosed in the context of a particular aspect, that feature may also be used in the context of other aspects to the greatest extent possible.

[0073] Furthermore, when a method having two or more defined steps or operations is mentioned in this application, these defined steps or operations may be performed in any order or simultaneously, unless the context precludes such possibilities.

[0074] All features disclosed in the specification (including the claims, abstract, and drawings) and all steps in any disclosed method or process may be combined in any combination, except that at least some of such features and / or steps are mutually exclusive combinations. Unless otherwise expressly stated, each feature disclosed in the specification (including the claims, abstract, and drawings) may be replaced by an alternative feature for the same, equivalent, or similar purpose.

[0075] It should be recognized that the above-disclosed and other features and functions, or variations thereof, can be combined into many other different systems or applications. Those skilled in the art can then make various substitutions, modifications, variations, or improvements therein that are not currently foreseen or anticipated, and these are intended to be included in the embodiments as well.

Claims

1. A magnetic resonance imaging transmitter coil, comprising: An input terminal, wherein the input terminal is used to receive one or more control signals; One or more voltage or current sources, said one or more voltage or current sources being used to generate one or more source signals; and At least two coil segments, each receiving one or more control signals and one or more source signals, each coil segment having one or more switches to convert the one or more source signals into one or more currents of arbitrary frequency, the one or more currents causing a magnetic field.

2. The transmitting coil according to claim 1, wherein, The coil includes any one of the following: surface coil, coil array, saddle coil, birdcage coil, solenoid, or gradient coil.

3. The transmitting coil according to claim 1, wherein, The one or more switches include one of an amplifier, a transistor, an integrated circuit, a diode, a relay, or a variable resistor.

4. The transmitting coil according to claim 1, wherein, The coil segment converts the one or more source signals into one or more arbitrary frequencies that excite the resonance of any one or more nuclei.

5. A method for operating a magnetic resonance imaging transmit coil, comprising: Generate one or more control signals; One or more source signals are generated using one or more voltage or current sources; Receive the one or more source signals and the one or more control signals at two or more coil segments; Switches are activated at two or more coil segments according to one or more control signals to convert the one or more source signals into one or more currents of arbitrary frequency to induce a magnetic field.

6. The method according to claim 5, wherein, The one or more control signals are configured to cause the one or more voltage or current sources to generate a current with a desired waveform and a desired duration.

7. The method according to claim 5, wherein, The one or more control signals are configured to cause the one or more voltage or current sources to generate currents at one or more desired frequencies to simultaneously or sequentially excite one or more cores.

8. The method according to claim 5, wherein, The one or more control signals are configured to cause the one or more voltage or current sources to generate currents for multiphoton excitation, spatial coding, or field homogenization.

9. The method according to claim 5, wherein, The one or more control signals are configured to cause the one or more voltage or current sources to simultaneously or sequentially generate currents at multiple desired frequencies for multiple purposes.

10. The method of claim 5, further comprising combining at least two coil segments with a receiving coil to simultaneously or sequentially generate an image of any nucleus.

11. The method of claim 5, further comprising combining at least two coil segments with a receiving coil to simultaneously or sequentially generate NMR spectra of arbitrary nuclei.

12. The method of claim 5 further comprises combining at least two coil segments with a receiving coil to generate an image or spatial location of an NMR spectrum, or an image or spatial location of a specific molecule discovered by NMR spectroscopy.

13. The method of claim 5, further comprising using the transmitting coil with one or any one MRI scanner having a variable static magnetic field, or with multiple MRI scanners having different static magnetic fields for the same or different nuclei.

14. The method of claim 5, further comprising using the transmitting coil to reduce the dead time between transmission and reception in the MRI system by providing shorter transmission pulses without resonant circuitry.

15. A magnetic resonance imaging (MRI) receiving coil, comprising: Two or more coil segments, each coil segment having: An input terminal that receives one or more control signals for the coil segment; A switch connected to the coil segment to change the available current path of the current induced by the received magnetic field; and One or more amplifiers, said amplifiers being used to receive power and convert said power into radio frequency (RF) input voltage signals of one or more desired frequencies.

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