DOI: 10.1148/rg.274065115
MR Imaging: Brief Overview and Emerging Applications1
Michael A. Jacobs, PhD,
Tamer S. Ibrahim, PhD, and
Ronald Ouwerkerk, PhD
1 From the Russell H. Morgan Department of Radiology and Radiological Science (M.A.J., R.O.) and Sidney Kimmel Comprehensive Cancer Center, Department of Oncology (M.A.J.), Johns Hopkins University School of Medicine, Traylor Bldg, Room 217, 712 Rutland Ave, Baltimore, MD 21205; the Departments of Radiology and Bioengineering, University of Pittsburgh, Pittsburgh, Pa (M.A.J., T.S.I.); and the School of Electrical and Computer Engineering and Bioengineering Center, University of Oklahoma, Norman, Okla (T.S.I.). From the AAPM/RSNA Physics Tutorial at the 2004 RSNA Annual Meeting. Received June 7, 2006; revision requested August 16; final revision received March 9, 2007; accepted March 9. Supported in part by grants 1R01CA100184 (M.A.J.), P50 CA103175 (M.A.J.), and 1R21CA095907-01 (R.O.) from the National Institutes of Health. All authors have no financial relationships to disclose.
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Figure 1. Typical gradient coil set used for localization of the MR signal. These coils are placed concentrically to each other within the magnet and are used sequentially for three-dimensional localization of the gradients to create images from the MR signal.
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Figure 2a. (a) Low- and high-flip-angle (arrows) images and measured transmit and receive fields obtained by using an 8-T system (top row) and corresponding simulated results obtained at 340 MHz by using computational electromagnetics (bottom row) (49). (b) In vivo 2000 x 2000 image of the human brain obtained at 8 T with 100-µm resolution (20). High-field-strength magnets are increasingly being used in MR imaging research centers throughout the world.
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Figure 2b. (a) Low- and high-flip-angle (arrows) images and measured transmit and receive fields obtained by using an 8-T system (top row) and corresponding simulated results obtained at 340 MHz by using computational electromagnetics (bottom row) (49). (b) In vivo 2000 x 2000 image of the human brain obtained at 8 T with 100-µm resolution (20). High-field-strength magnets are increasingly being used in MR imaging research centers throughout the world.
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Figure 4. Readout gradient for three discrete frequencies associated with three positions and their summed signal. Also shown are the actual readout signal, which is the sum of signals at all frequencies within the bandwidth (BW), and the fast Fourier transform (FFT) of this signal, which is a projection of the object along the frequency-encoding axis.
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Figure 5. Phase encoding. The readout experiment shown in Figure 4 is repeated N times (N is the desired image resolution) with a short gradient pulse of amplitude GPE(k) and length tPE preceding readout. This gradient pulse temporarily changes the frequency; after period tPE, the result is a phase shift  (k,r). If the gradient pulse amplitude GPE(k) or length is varied in N equal steps, the resulting set of phase-encoded profiles (after the fast Fourier transform of the readout direction) will be the Fourier transform of the object along the phase-encoding direction.
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Figure 7. T1-weighted images (T1WI), T2-weighted images (T2WI), and diffusion-weighted images (DWI) from different regions of the body with corresponding maps of the apparent diffusion coefficient (ADC) of water. Top: Brain images of a patient with an acute stroke (<6 hours) and older infarct (>3 months). The regions of infarction are clearly visualized on the T1- and T2-weighted images as low (T1) and high (T2) signal intensity in the left temporal lobe (open arrow). Conversely, in the right occipital lobe, there is little or no change in the regions of new ischemia, except that they are seen as hyperintense areas on the diffusion-weighted image (single solid arrow). On the ADC map, there are corresponding hypointense regions (double solid arrow), which have lower ADC values. Areas that are hyperintense on the ADC map have higher ADC values. Similar signal intensities are noted on images of the breast (middle) and uterus (bottom). Note the changes (arrow in bottom row) on the diffusion-weighted image of the uterus, with decreased signal intensity in the same regions on the ADC map. Similar changes are seen on the breast images (arrow in middle row). These examples demonstrate the versatility of MR imaging in producing different image contrasts.
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Figure 8. Effects of removing k-space data on a reconstructed phantom image. In A, image was obtained with full k-space. In B, image obtained with the center of k-space missing shows only edges and fine detail, which are defined by high-frequency k-space. In general, the maximum signal is obtained from the center of k-space. In C, image obtained with only the outer portion of central k-space removed is blurred and lacks detail. In B and C, note that removal of portions of k-space leads to ringing artifacts in the image (56).
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Figure 9. Path through k-space of a gradient-echo sequence with phase encoding. Sequence diagram shows the excitation pulse and signals RF, the readout gradient GR, and the phase-encoding gradient GP. The prewinding gradient A carries the k-space trajectory in the kx direction out of the sampling area. Phase encoding with B causes an offset of the trajectory in the ky direction, where signal from one k-line is read out under C. Data are acquired only during the last part (solid line in the trajectory), and signal during A and B is discarded (dashed line). The experiment is repeated with different B until all k-lines have been acquired (58).
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Copyright © 2007 by the Radiological Society of North America.
) for localization of the MR signal. RES = resolution. 2, Use of the RF pulse to invert the magnetization signal within the voxel (
= time of the RF pulse and area covered by the pulse (eg, the strength of the pulse). M = magnetization. 3, Graphic representation of the relaxation parameter T1. TR = repetition time. 4, Graphic representation of the relaxation parameter T2. TE = echo time.