Prototype halbach-bulb scanner is shown. The order of the components from innermost to outermost are: RF Tx/Rx coil, B0 shim magnets, RF shield, main B0 magnet, and gradient coils. During the scan, the patient’s head rest on the RF coil. The rest of the system (B0 magnet, RF shield, and gradient coils) is on rails and slides over the coil and patient.

3D T2-weighted images of the brain in a healthy adult volunteer. 3D RARE sequence, resolution = 1.5mm x 2.8mm x 7mm,TR/TEeff =3000ms/134ms, acquisition time = 25 min (6 averages). A) Images reconstructed assuming linear encoding fields resulting in distortion, b) Images reconstructed using the generalized encoding method employing the measured fieldmaps in (c). C) combined measured B0 variation map and readout gradient maps for image reconstruction.

(a) Picture of the scanned rabbit head. (b) Picture of a rabbit skull, taken from Gabrielle Ochnik, Pinterest. (c) Top: single slices for 90 µs dead time; middle: the same slices for 1 ms dead time; bottom: difference between the above images.

(a) 2 dimensional slices of four human teeth embedded in a piece of pork ham (PETRA). (b) Photograph of the sample.

Figure 1: (a) Proposed C7T magnet showing the 8-coil configuration, cold heads, and helium tanks. The size of the C7T relative to an average person is also shown for scale. (b) One of the 2 large main coils being wound with superconducting NbTi wire.

Figure 4: Diffusion imaging performance at different b-values showing (a) diffusion TE time, and (b) diffusion pulse width (d) for gradient configurations of whole-body 7T, C3T, C7T, and MAGNUS. The performance improvement between the C7T gradient and the whole-body 7T is clear.

Figure 2 - Left) Photo of gradient coil with 3 stepped layers. The overall mechanical length of the gradient coil is 160 cm to mount easily inside magnet and for better dynamics. The total extent of the wires in Z direction is 119 cm. Right) Plot of wiring pattern showing 3 layers as primary (Pri), intermediate (Mid) and secondary (Sec) layers in stepped design.

Figure 1 - Photos of the scanner. A) Front side of scanner showing RF cables connecting the 96-ch Rx, 16-ch Tx coil to 128-ch interface box. B) gradient coil cover with shoulder cut outs with the coil interface box in front. The front bore is 60 cm diameter for subject access. C) Photo of energy chain between two 64-ch RF receiver interfaces. D) The rear bore is 39 cm diameter. The coil interface box is positioned in the rear of the magnet bore when the RF coil is positioned at isocenter. The flat flexible energy chain shields over 270 cables and is seen connected to the coil interface box.

Figure 1: The components of the setup used. The NG500 1.1 gradient amplifier (A). (B) consists of the 2-axis insert gradient, an 8 channel transceive dipole array, a 64 channel receive array and an agar-based head phantom. The middle images show a more detailed view without insulation tape of the used insert gradient and receiver coils.

Figure 4: Effect of the oscillating gradient on the MR-signal. Left: the MR-signal with the amplifier off. Middle: the amplifier on. This data was acquired with a slice-selective 1D acquisition showing only one of the simultaneously acquired 72 channels. Right: raw data controlling all 5 gradients.

Figure 3: Single- and two-photon GRE brain images of a healthy volunteer. Left panel: Single-photon image with a flip angle of 44.5 degrees vs. a TR of 350 ms (Top) and a TR of 150 ms (Bottom). Right panel: Two-photon image with a $$$B_{1xy}$$$ amplitude corresponding to an equivalent two-photon 44.5-degree flip vs. a TR of 350 ms (Top) and a TR of 150 ms (Bottom). Some brain structure contrast is seen to be enhanced. Other parameters: FOV 27.6 cm, slice thickness 3 mm, matrix 256x160, TE 10ms.

Figure 2: Left: equipment room set up. Right: $$$B_{1z}$$$ coil in a 3T scanner.

Figure 4: Top: Number of reported stimulations in different body parts during the experimental study. Bottom: Maps of predicted PNS hot-spots in the male and female model given in terms of PNS oracle and equivalent PNS thresholds (inverse of the PNS oracle) for a 400 us rise time trapezoidal waveform. Every blob corresponds to an activation hot-spot, with both color and size corresponding to the strength of the activation.

Figure 1: Photo of the final constructed Impulse gradient coil and 3D rendering of the three-layer winding pattern (primary, intermediate, and secondary layers of all axes combined). The gray sphere corresponds to the 20 cm region-of-linearity (ROL). Note that the constructed coil is longer (160 cm) than the wire extent (119 cm) to improve mechanical properties and simplify the MR system assembly.

Figure 4: In vivo respiratory data from the wireless accelerometer (blue) and video (orange) show excellent agreement during normal breathing (~0.25Hz, R>0.9) and lack of interference from MRI through closely matching waveforms and derived respiratory frequencies. The accelerometer and video had slightly lower correlation (R>0.86) during heavy breathing (~0.8Hz), possibly because the respiratory frequency approached the sensor’s current sampling frequency (2 Hz). Note the video (60 fps) was resampled to match the time basis of the accelerometer to display their correlation.

Figure 2: Left panel: Photos alongside frequency maps show devices that are compatible or incompatible with local use because field disturbance is limited (a,b,c,e) or not limited (d,f) to voxels within 2cm of the surface, respectively. Right panel: Noise spectra using the ATWINC1500 show similar characteristics at baseline and with the shielded wireless device for 23.5±0.25MHz, but show interference at 24MHz (likely caused by a harmonic from the 2.4GHz WiFi signal).

Figure 3: Retrospective motion correction of a structural scan with instructed movement (Table 1: subject 2, trial 4) with moderate motion of 1-2mm (x,y,z) and 1-2deg (θ,ϕ,ψ). Tracking curves for CWLs and NAVs (red, blue respectively) show high degree of similarity (low mean absolute difference, MAD), with largest discrepancy (gray) in y-translations. A substantial increase to visual sharpness (e-l) is evident after retrospective correction using either tracking method. Both methods lead to an increase to average edge strength (AES) of 12.8/9.9% for CWLs and NAVs, respectively.

Figure 1: Measured wire loop signal $$$\mathit{v}_i(t)$$$ for part of a navigator consisting of a weighted sum of columns of $$$\dot{\tilde{\pmb{\mathit{G}}}}(t)$$$, which resemble filtered time derivatives of gradient activity, and are found by sampling the sequence in a static phantom pre-scan with only one active gradient component $$$\mathit{g}_x(t)$$$, $$$\mathit{g}_y(t)$$$, $$$\mathit{g}_z(t)$$$, or active RF. Each wire loop ($$$i=[1,2,\ldots ,I]$$$) measures a unique weighted sum dependent on loop position, orientation, and geometry.

Figure 1: A schematic illustration of (a) Beat Pilot Tone (BPT) vs. (b) Pilot Tone (PT) acquisitions. PT has a 2.347m wavelength ($$$f_{PT} = 127.8MHz$$$), while BPT has a 12.5cm wavelength ($$$ f_T = 2.4GHz $$$); thus, there is much more interaction between BPT and the cm-scale motion of the body. c) BPT uses two frequencies, $$$f_T$$$ and $$$f_T+f_{BPT}$$$, where $$$f_{BPT}=f_{larmor} + f_{offset}$$$ (e.g., $$$f_{offset} = 100kHz$$$). A signal with a beat frequency $$$f_{BPT}$$$ is mixed in the coil preamplifier due to intermodulation; this signal is the BPT.

Figure 4: a) A volunteer was asked to move their head up and down during the scan, a motion in the range of 2cm. b) BPT from the 3 most modulated coils are compared to the PT, with the coil arrangement in c). The arrows show up-down motion, which is clearly visible in BPT, but barely so in PT. d) The two most modulated body coils (17 and 20) for the abdominal scan. e) The BPT and PT magnitudes and phases were filtered with a moving average and scaled to match the bellows signal, which they match closely.

Figure 2. Magnitude (A), time-domain (B) and phase (C) of measured gradient IRF for the three axes, without pre-emphasis compensation. The arrows highlight channel specific known mechanical resonances at lower frequencies. No filtering has been used on this data. Notably, measurement noise in the IRF increases towards the higher frequencies, where the power of the input waveform is null.

Figure 4. Representative images for two different readouts for a 2D spiral acquisition. For the shorter readout (A) both monitored and IRF demonstrate distinct sharpening of the grid lines and decreased blurring at the edges, compared to nominal reconstruction, also evident in the difference images (B). For the longer readout (C), error in correction for the field monitored measurement is a result of the probes dephasing during the long readout (D). IRF predicted trajectory correction can be used to compensate for this.

Figure 4: Phantom UTE images reconstructed with GIRF/delay correction a) off/off; b) off/on; c) on/off; d) on/on. Note artifacts: halo (solid arrows), dark boundaries (dotted arrows), and bright boundaries (dashed arrows).

Figure 3: First-order GIRFs in frequency space measured for the MR-Linac at our institution. Represented in terms of magnitude (arbitrary units) and complex phase components for physical X, Y, and Z gradients of the system.

Figure 2: Magnitude of the linear self-term of the GSTF in x-direction, evaluated with three different datasets: All 18 readouts from all six measurements (top), only the first readout of each of the six measurements (middle) and only the first readout of the first measurement (bottom). For better visibility, the curves are plotted with an offset of 0.05 between each other.

Figure 1: Schematic sequence timing of the first and second measurement. The durations and amplitudes of the RF pulses, gradients and readouts are not displayed to scale. For clarity, only three of the 16 triangles are depicted. Each readout had a duration of approximately 25 ms.

Figure 2: Phase error from a single-sided encoding experiment for (a) MAGNUS unit#2, and (b) the C3T system. The difference between the measured and the EM-derived nominal offset of 12 cm is shown.

Figure 4: Phase difference images for the single-sided encoding experiment with g₃ = 71 mT/m (a) before applying the zeroth order frequency correction (ROI phase = 4.2$$$\pm$$$0.1 radians), and (b) after correction (ROI phase = 0.1$$$\pm$$$0.1 radians). A zero phase is expected for the phase difference in the stationary phantom.

Figure 2: Comparison of Gradient Modulation Transfer Functions’ (GMTF) self terms for 0.75T and 3T configurations. Mechanical resonances are identified by local dips (anti-resonances) and peaks (resonances) in the GMTF that deviate from the ideal low-pass GMTF. Resonances are generally less pronounced for 0.75T compared to 3T. A reduction by 60% for 0.75T is found for the antiresonance in the z axis GMTF.

Figure 4: Comparison of relative microphone amplitude spectra (RMAS) for the chirp sequence at 0.75T and 3T. For each gradient axis, RMAS were normalized to the maximal value. RMAS are reduced for 0.75T by up to 70% depending on the gradient axis compared to 3T. In addition, peaks in the RMAS correspond to the mechanical resonances identified in the GMTF.

Figure 3: Reconstruction results without and with inclusion of the PSF-mapping (top row). Here, the zoomed FOV is shown (112 x 112 mm2 to show more detail in the phantom). The differences between the different reconstructions (bottom row). Note that we can clearly see the large effect of the 0th order PSF component on the ghosting. The horizontal lines in all images are caused by a stimulated echo.

Figure 4: Reconstruction results for the 0th order correction on a large head phantom. Due to the large phantom size, large B1+-inhomogeneities are visible in the bottom of the phantom. In addition, some air bubble induced susceptibility artifacts and signal pile-up due to field non-linearities are visible in the bottom of the image

EPI reconstructed with concurrent field monitoring (without SENSE and without B0 correction). Left: reconstructed using corrected probe signal for the trajectory calculation, middle: reconstructed without probe signal correction, right: difference image. The ghosts disappear when the concurrently monitored probe phases, used to calculate the trajectory, are corrected.

Phase correction of an x-gradient frequency swept pulse (20 mT/m). In the zoom it can be seen that the corrected signal (red) coincides with the signal that is measured in absence of the head coil (blue).

Figure 4. (a) Measured trapezoidal current waveform with 50 A flat top amplitude and 0.25 A/us slew-rate with filter. (b) Temperature of eGaN transistors to 45° C from room temperature for the current on (a) with 10% duty cycle and five minutes of operation. Zoomed version of the current is shown in Fig. 5.

Figure 5. (a) Zoomed current (top) and voltage (bottom) applied to the coil in Fig. 4 without (a) and with (b) filter. Peak-to-peak ripple current is reduced more than twelve times by using the LC filter. Small resistance of the coil (about 200 mΩ) requires small voltage levels and PWM duty cycle.

Fig.4. Experiment results. Comparison of the coils’ currents for the linear (uncompensated) and nonlinear (compensated) models. The reference inputs are trapezoid waveforms with 50A and 10A at the flat-top for ch-1 and ch-2, respectively. The rise time is the same for both channels (200μs).

Fig.1. (a) The circuit diagram of the two-channels gradient array. (b) The hardware setup including control card (VC707 evaluation board), custom-built GPAs and two channels array coils.

Figure 2: (A) Means (SD) of reported sound level ratings directly after the scan (immediate) and after all scans (delayed) for the quiet compared to the standard scan; (B) Means (SD) of comfort level, overall experience and willingness to undergo scan again ratings for the quiet compared to the standard scan. Asterisk indicates statistical significance.

Figure 3: T₁-weighted axial images of the standard (top row) and the quiet (bottom row) scans of all five subjects.

Figure 3. The simulated electric field on the yz plane at the isocenter within the x and y axes coils with 1A of current flowing through the coil, a. showing the contribution from the x axis gradient coil and b. showing the contribution from the y axis coil.

Figure 2. The simulated electric field on the xy plane at the isocenter of both the x and y coil along a specific line with a. evaluated at the point shown in Figure 1a. and b. evaluated at the point shown in Figure 1b. with 1A of current flowing through the coil.

Fig 4: Diffusion-Weighted images (top row) measured in a dedicated phantom within the same slice for different amplitudes demonstrate the diffusion encoding capability. Measurements were performed with spin-echo diffusion deploying 10ms duration for each gradient pulse. ADC maps (bottom row) calculated using measured field maps to correct for inhomogeneities. Depicted are ADC maps acquired with same currents which results in different diffusion weighting closer to the bottom of the cup. Positions are given in the frame of reference of the scanner.

Fig 5: Breast gradient coil during construction. Coil windings from square profile litz wire were manually wound onto a 3D-printed former (upper left). All components were combined into a housing (right) including water cooling, temperature sensors, and a second, non-functional cup. The coil was cast in epoxy (lower left) under vacuum.

Figure 2: Coil track location of the upper disc of the two-channel coil.

Figure 5: Dissipated power of two-channel coil designs in comparison to the conventional coil design.

Fig. 1 The 3D model diagram of the proposed biplanar matrix coil and the open-structure low-field MRI system. The green and red boxes indicate the specific installation location of the coil. The red and green arrows indicate the biplanar matrix coil and the low-field MRI system, respectively.

Fig. 4 Contour and field of the proposed biplanar Matrix Z2 Nonlinear Gradient Coil (a) Matrix coil composed of coil elements with the same structure. The green and red boxes indicate the position of the coil, as seen as Fig. 1. (b) The model diagram of the different CP’s positions of the Z2 spherical harmonic field being moved at different times in (a). The energization sequence of the biplanar matrix coil of 4×4 element is from t8 to t9, and t7 to t10 follow the arrow direction.

Figure 1: A quarter geometry of the proposed gradient system. Z-gradient array consists of 12 pairs of wire bundles, each with 10 copper wires of 2 mm diameter. The diameter/height of the main and shield array coils are 24/30 and 30/35 cm, respectively. Copper sheets of 2 mm thickness and 40 cm height for the main coil and square copper wires (2x2 mm²) are used for shield coils. The diameter of the main/shield coil for the x- and y-gradient coils are 20/26 and 22/28 cm, respectively. The aluminum shell of 90 cm diameter (the warm cryostat) is not shown here.

Figure 4: The same gradient system with the second configuration for the z-gradient array profile. In this case, the FOV size is reduced to 60 mm and it is shifted up 30 mm along the z-axis. Z-gradient strength is doubled to be 250 mT/m within the shifted FOV region with less than 8.9% deviation from linearity. The contour lines indicate 1 mT separations. The residual eddy current is 0.007%.

Figure 4. (a) Diagram of spokes-and-hub magnet with no tilt (blue outline) and a 0.25° tilt (orange outline) about the x-axis, creating a y-axis gradient field; (b) Simulated receive signal spectrum with and without tilt; (c) Measured spin echo obtained from 8mm diameter test tube of water using dithered ultrasound RF pulse with 100 kHz bandwidth with 0° (blue) and 0.25° (orange) tilted magnet; (d) Measured received signal spectrum with (orange) and without (blue) tilt.

Figure 5. Repeat of experiment in Figure 4, but with an 8mm diameter phantom containing two distinct tubes of water. In (a), tubes are side-by-side (along the y axis and in line with the tilt-induced gradient field); and in (b), tubes are stacked (along the z axis, orthogonal to the gradient field). Note the clear double hump in (a) showing spatial separation.

The illustration of (a) Nested for-loop, and (b) Approach-1 and (c) Approach-2 for acceleration of calculation of the magnetic fields of PMAs.

The configuration of (a) Halbach (60 blocks) [6] and (b) IO ring-pair PMA (1200 blocks) [5]

Theoretical calculations in 2D consisting of four coils (a) and comparison of magnetic field values with experimental results from A to B (b).

3D representation (a) and measurement points (b) of constructed system

Fig. 5 First column: the unimodal distribution of the 0.5mmol/L MnCl₂ solution. Diffusion and relaxation coefficients measured by CPMG-only are identical to the results of SE-CPMG. Second column: D-T₂ distribution of the peanut oil. Two different fat components are demonstrated. However the distributions, especially for the T₂ values are different. Third column: D-T₂ distribution for the combination of 0.5mmol/L MnCl₂ solution and peanut oil. Despite the differences in the D-T₂ distributions, both (c) and (f) show clear separation of fat and water components (three peaks).

Fig. 2 Details of the single-sided NMR experimental device. (a) The arched magnet with the RF coil in the magnet center. The geometry is designed to position the device at the waist abdomen for liver detection. (b) Two different plastic bottles are placed inside the magnet, filled with peanut oil and 0.5 mmol/L MnCl₂ solution, respectively. The experiment is designed to explore the sequences for separate characterization of oil and water simultaneously.

Figure 5: In-vivo B₀ field mapping of the liver in a healthy volunteer during free breathing with two different B₀ shimming methods: 1) scanner shim (2^nd order) and 2) array of shim coils in combination with the conventional second scanner shim. A: anterior, P: posterior, R: right, L: left, H: head, F: feet, and std: standard deviation.

Figure 1: A) The array of shim coils B) used around the body array in the 7T MR system

Figure 1 Circuit diagrams of RF-only coil, RF coil + DC coil, and RF coil + “RF transparent” DC coil. Note that in practice, feed wires of DC coils are typically twisted to cancel the Lorentz force. For simplicity, straight feed wires were shown here. For Bazooka-feedline [5] few-turn “RF transparent” DC coil, the shield of twisted wires (could be one shield for each wire or one shield for both wires) was shorted at the position that is 1/4 wavelength away from the DC coil, forming open circuits and block RF signal going through.

Figure 2 Circuit diagrams of many-turn “RF transparent” DC coil using float trap design. For many-turn DC coil, it may resonate close to the RF frequency even the feed port is broken using the bazooka feedline method (A). In this case, it still can be seen as a coupled lossy resonator and leads to SNR loss. The float trap [6] is a concentric resonator that inductively couples and traps the RF signal going through it. It is expected to break every turn of the DC coil that goes through the trap and thus avoid unwanted resonate modes.

Figure 1: a) From left to right: Ideal channel disposition and geometry for first and second layers, both layers resulting from the SVD-based MCA optimization; and ideal 2-layer 36-channel system. b) The two prototype layers and the assembled shimming system with RF housing in the interior of the shim layers. c) Simulation models of generic 24-ch (left) and 48-ch (right) M-MCA used for comparison. Circular coil radii are 4.5-cm and 3.5-cm, respectively.

Figure 5: a) Brain mask overlay (green) considered for optimal current calculations for global shimming and selected slices (blue) presented in b) rs-EPI acquisitions. Baseline, global and slicewise (dynamic) 2-layer 36-channel SCOTCH shimming are shown. Gradient Echo rS-EPI sequence parameters were 2-mm in-plane resolution, fifteen 3-mm-thick slices, TE/TR=44/84000ms. Yellow arrows highlighting zones with significant improvement.

Figure 1 Three RF loops (84mm diameter each) were arranged in the configuration shown on the left. The four different settings for the shim coildiameter/positioning used in the numerical simulation are also shown. The reference S-parameter curves (in the absence of the shim coils) are shown on theright.

Figure 4 Optimal arrangement of 24 shim coils on a cylindrical holder (units in [m]).

Figure 3 – The 6 channel transceive array facilitates good coverage over the cervical spine (a), although there is a large B₀ background field present (b). The transceive array with B₀ shim capabilities can provide additional B₀ fields (c, sagittal). The six channels show non-overlapping RF fields (d) as well as complementary B₀ fields (e) even for 1A current.

Figure 1 – Extension of RF coils with DC circuits. In a traditional loop coil (a) the conductor is broken up with capacitors whereas (b) shielded coaxial coils (SSC) do not require this. Adding a DC path (red) therefore typically requires the addition of several RF chokes in a traditional coil (c) compared to only two on an SSC coil (d). Addition of RF chokes is only required on the feeding points (e) and not limited additional space is needed on the coil array (f). Note the coil array shows virtually no coupling between the elements despite lack of overlap between elements (data not shown).

Figure 3. Phantom axial GRE images of cross sections at the femoral head (a,b) and near the femoral insertion end (c,d) of the metal hip prosthesis at TE=4.80ms before (left) and after (right) adding UNIC shim to scanner shim. Signal void is stronger at the longer echo time due to additional dephasing between the two echos. UNIC shimming reduced signal void artifact area by 9% (b versus a) and 9% (d versus c), and extended the visible area closer to the metal, increasing it by 50% (b versus a); and 8 % (d versus c); see text for ROI definition. Fiducial markers are visible in a and b.

Figure 5. B0 field homogeneity improvement with UNIC shimming. Histograms of Larmor frequency deviation from the scanner center frequency (~123MHz), of the masked volume of pixels in the 3D slab scanned before and after adding UNIC shim to scanner shim (a), and B0 field magnitude maps (b) for the slice in Figures 3c,d before (b. left) and after (b. right) adding UNIC shim to scanner shim.

Figure 2: Realtime z-shimming overview. Training session: Step 1: The pressure and fieldmaps are acquired. Step 2: Shimming Toolbox’s optimizer computes a time series of Gz maps. A linear regression is computed between the field gradient and the pressure. Imaging session: Step 3: A dummy GRE scan is run in order to obtain the target image. Step 4: The $$$G_z$$$ maps are resampled onto the target GRE images and $$$G_{z, static}$$$ and c for each slice are sent to the scanner. Step 5: Run the custom realtime z-shimming GRE sequence.

Figure 1: Pneumatic phantom with a long stick, which terminates by a slightly ferromagnetic object that oscillates back and forth creating a time-varying magnetic field.

Figure 1 (a) Schematic diagram cylindrical shield with end-caps of radius, b, and length, 2Ls, and coil of radius, a. Illustrative coil wirepath and direction of current flow indicated. (b) Multiple reflections of the coil due to the end-caps (dotted lines indicate reflected versions of the wirepath).

Figure 2 (a) Experimental set-up: fields from loop and saddle coils measured using a fluxgate magnetometer inside a 30 cm diameter 1 m long cylinder formed from 1.5 mm thick mu-metal. (b) Helmholtz-type coil formed from two loops of radius a, separated by a distance 2d, carrying current in the same sense.

Figure 3 – Phase variation during 2H MRS interleaved with a 7-slice spin-echo MRI. The phantom consisted of three bottles containing ~0.1% D2O and various amounts of DMSO-D6 (~0.02 – 0.05%). A small (0.5 mL) phantom containing pure D2O served as an external off-resonance reference signal. (A) Phase variation in the absence of a phase lock. (B) In addition to the strong linear phase roll there is also a smaller sequence-dependent phase modulation. (C) In the presence of a phase lock, each spectrum has an identical phase. The 1D spectra shown in (A) and (C) are extracted from repetition 200.

Figure 1 – (A) Four channel interleaved mixer. Inputs for the mixer are from their respective preamplifiers. The 1H signal is applied to one arm of the RF switch. The 2H signal is sent to an RF mixer and upconverted to 170MHz by mixing with the 144MHz LO. The upconverted signal is filtered to remove the unwanted sideband, amplified and sent to the RF switch. The PPG controls when each nucleus is sent to the scanner’s 1H receiver. (B) Each channel is isolated from outside interference and each other by using a dedicated RF shielded enclosure built onto the board.

Figure 1. (a) The open MRI system. (b) The diagram of simultaneous transmit and receive. The phantom is above the coils. The coil on the right side is for the use of picking up NMR signal at 1MHz, and the coil on the right side is irradiating offset field.

Figure 5. Experiments for simultaneous transmit and receive. (a) The diagram of a multiple spin echo sequence with continuous offset irradiation in the front five acquisition window. (b) Experimental data acquired with the sequence above with echo spacing (TE) 1ms. It shows three echo trains under different offset-field irradiations.

Figure 4. Measured actual pulses (vertical dashed lines indicate the input pulse ends.) of the sinc pulses (top row) and compensated sinc pulses calculated with N_S values of 13 (second row), 25 (third row), 37 (bottom row) and bandwidths of 30 kHz (left column), 90 kHz (middle column), and 150 kHz (right column).

Figure 2. (a) Schematic diagram of an open, table-top, MRI system with photos of (b) the RF solenoid coil and (c) the magnet. (d) A spin-echo sequence with pre-polarization.

Figure 1. (a) Scheme and (b) photo of the 3-axis probe. The probe consists with three perpendicular square loops.

Figure 2. Experimental set up.

Figure 2. (LEFT) Percent difference in dreMR image of VivoTrax, in uniform concentration of 160 μM, using field of previously constructed coil. Percent difference is taken from center point where field is at desired value. (RIGHT) Same simulation using field from new design method coil. Colour axis chosen to match the left.

Figure 1. (LEFT) Simulated dreMR image of 4 cylinders of VivoTrax, in concentrations of 20, 80, 120, and 160 μM, using field of previously constructed coil. Note change in signal across cylinders. Sequence uses 300mT field shift in a 0.5T system. (RIGHT) Same simulation using field from new design method coil. Signal does not noticeably change across cylinders. Data points are given in both figures at locations with no contrast agent.

Figure 3: The very low field MRI in the magnetically shielded room.

Figure 1: The eight channels sensor. Each channel is matched and tuned, and decoupled from emission radiofrequency.

Measurement setup on the Tabletop-system. Left side: OCRA-console for controlling the system, supplying the components and connecting to the computer. Right side: Opened system with the sample in the carrier between the neodymium magnets with connected signal and shimming lines.

Measurement setup on the Skyra-system. The test tube sample was positioned with the coil perpendicular to the static magnetic field fixed by wooden planks on the patient table. The connection to the MR system was established via the T/R-switch and TIM-connector. Shown on the bottom is an image taken with the patient camera.

Figure 2 Abdominal imaging at 5T. a) and c) B1+ sensitivity map and T2W FSE image with CP mode excitation. b) and d) B1+sensitivity map and T2W FSE image after static RF shimming optimization. The CV values were measured for bothcases.The red arrows show the most significant improvement area.

Figure 3 Pelvic imaging at 5T. a) and c) B1+ sensitivity map and T2W FSE image with CP mode excitation. b) and d) B1+sensitivity map and T2W FSE image after static RF shimming optimization. The CV values were measured for bothcases.The red arrows show the most significant improvement area.

Cooling system schematic. A: Pre-cooling / setup configuration - The chiller is operating and pre-cooling the bulk of the fluid to the desired temperature, while the cooling pad and temperature sensor(s) are being used for sample preparation. B: Normal operation configuration - The temperature probe(s) are affixed to the sample, and the cooling pad is wrapped around it. A layer of cloth around the outside provides thermal insulation to prevent excessive condensation, and the pad is connected to the chiller.

Example T2- and diffusion weighted images. Top: T2 weighted images were acquired at 0.4mm isotropic resolution using a 3D-SPACE sequence with FOV=150x150x96mm3, TR/TE=2300/586ms, 6 repetitions. Bottom: averaged diffusion weighting images acquired at 0.8mm isotropic resolution using a spin-echo diffusion weighted readout-segmented EPI sequence with FOV=150x150x104mm3, TR/TE=10.8s/113ms, multiband factor 2, GRAPPA factor 2, b=9000 s/mm2.

Figure 1: Design and conception of the bioreactor. (A) 3D model of the bioreactor and its support. (B) Detailed view of the bottom of the 3D coil circuit and its integration. (C) Detailed view of the top of the 3D coil circuit and its integration

Figure 3: (A) Picture of the bioprinted tissue after construction. (B) 2D TURBO RARE image of the tissue using the bioreactor acquired on a 7T MRI scanner with a 75 µm in plane resolution.

Figure 3. Synchrotron micro-CT scan data at two zoom levels, transformed to align lines of material with the viewing planes. Upper row: View of the entire scan ROI. Lower row: Detailed view of a short length of five stacked lines of material. The direction of print-head motion was left-right in the left- and right-most columns, and perpendicular to the image in the center column.

Figure 1. a) Confocal microscopy z-stack image of a stained phantom slide. Elastomeric matrix (red) and pores (black) are visible. Each white outline indicates an individual line of material. b) 2D projection of a 3D microscopy volume acquired with confocal microscopy. Outlined in yellow are larger pores caused by the printing pattern of the phantom. In both a and b, the image plane is perpendicular to the long axis of the pores.

Fig. 1: Schematic representation of the MLV preparation method. (a) Soy PC, alone or as a mixture, was dissolved in chloroform (b) put in a vacuum rotary evaporator to create a uniform lipid film. (c) Ammonium bicarbonate was added to dissolve the lipid film. (d) The solution was then lyophilized and (e) rehydrated with a desired iron ion buffer to create the iron ion encapsulated MLV. (f) The samples were imaged in a clinical scanner.

Fig. 3: qMRI relaxation rate R2* with increasing concentration of iron ion Fe²⁺. R2* is affected by the lipid-water fraction and iron concentration. R2* values increase with increasing concentration of Fe²⁺ for all lipid types. Each color represents a different lipid-water fraction MLV phantom. (a) Lipid MLV phantom PC-Cholesterol with Fe²⁺. (b) Lipid MLV phantom PC with Fe²⁺. (c) Lipid MLV phantom PC-SM with Fe²⁺.