MPR Image Quality

The quality and utility of MPR images is directly related to the quality of the base images used to create them. In this sense the factors affecting image quality are the same as for any scan. Suitable contrast, SNR, contrast to noise ratio (CNR) and the minimisation of artefacts are most important. Because most MPR work involves the use of short TR GRE 3DFT sequences, obtaining suitable image contrast can be difficult. As will be discussed later, the best solutions are offered by complex and efficient GRE sequences. 
Geometric parameters assume an added importance in this application. Slice thickness cannot be used to provide SNR to compensate for losses inherent in providing high in-plane resolution. Optimum results require slice thickness equal to the in-plane pixel size (isotropic data). When the data is not isotropic, images reformatted using the larger dimension will show reduced spatial resolution. This problem is worst when the large dimension is in the reconstruction plane, but its visibility will depend on the level of anatomical detail that needs to be resolved. 
Artefacts causing geometric distortion of the base image will also degrade the MPR image fidelity. Contrast and signal loss from cross talk becomes a problem when designing thin slice contiguous 2DFT sequences, particularly in multi-slice sequences with low acquisition bandwidths. While attempting to maintain a reasonable number of slices per unit TR, the sequence designer may choose a shorter period RF pulse, to make up for the increased sampling time dictated by a low acquisition bandwidth, sacrificing pulse shaping and slice profile in the process. 
The MPR software will employ specific algorithms to assign the pixel values of a reformatted slice. The image quality will depend on how well these handle multiple voxel values in thick slices, and how the data gaps resulting from inter-slice gaps are interpolated. The interpolation routines of magnification programmes can also significantly affect presented image quality. 
In any given system best results are obtained with small dimension isotropic data sets acquired from 3DFT data. When using 2DFT sequences, keep slice thickness and slice gap as small as possible, and plan to reconstruct planes at small angles to the acquired plane.

For best results use (in order of preference):

High SNR sequences with a: -
3DFT small dimension isotropic data 
3DFT small dimension anisotropic data 
2DFT thin slices with minimal or no inter-slice gap and optimised slice profile.

Artefacts in 3DFT

Any artefact that can affect a 2D image can affect a 3D image in the same way, although they can appear differently due to the difference in spatial localisation method. The phase encoding for partitions is prone to the same motion ghosts as normal in-plane phase encoding, but they will extend in the slice select direction and be noted on adjacent images. These can be seen across the slice select direction so the ghosts of the pulsatile object may not appear in the same slice as the artefact, nor in every slice. In multi-slab 3DFT, ghosts in the slice select direction will be restricted to the slab containing the source. 
Aliasing is common in the slice select direction whenever the object is larger than the slab width. The signals from outside one edge of the slab will wrap around to the slices at the opposite edge of the slab. This is commonly seen as extra ears on sagittal head slabs, or bilateral display of the fibulae in knee sequences. Restricted view coils can minimise this aliasing. Phase over-sampling in the slice select direction is more effective but costs time. 
The profile of the slab is prone to the same distortions as 2DFT slice profiles, and the poor slice profile common to many short TR selective 3DFT sequences is frequently seen. Edge slices will appear with poor signal level and little contrast. The severity of the effect depends entirely on sequence design and needs to be assessed individually. Its appearance is often compounded by slice direction aliasing. It is overcome by extending the thickness of the slab by 10-30%, acquiring extra partitions and discarding the poor images. 
The slice profile of a partition is rectangular, so there is no loss of contrast due to cross excitation (cross-talk) between partitions. 

Techniquefor MRI Liver to find out small HCC

No single MRI technique is optimal for imaging the cirrhotic liver, and individual preferences are usually based on the hardware and software available. We strongly encourage use of a body phased-array coil for the best signal-to-noise ratio. We prefer a 1 to 1.5-tesla magnet with a gradient rise time of at least 600 msec.

T2-weighted (or STIR, for short tau-inversion recovery) images are an essential component of the MR exam. While conventional spin-echo pulse sequences provide excellent image contrast, long acquisition times may result in substantial motion artifact. Many centers have abandoned this technique for faster echo-train fast spin-echo pulse sequences.4 These may be performed with respiratory gating or respiratory ordered phase-encoding, depending on the manufacturer.

By decreasing the repetition time (TR) and using a relatively long echo-train, T2-weighted images can be acquired in the time frame of a single breath-hold. Frequency-selective fat suppression may be used to augment image contrast, which is especially helpful if the liver is fatty. Alternatively, fat-suppressed images with additive T1 and T2 contrast can be performed with STIR methods when inversion times are selected to null fat. This technique is less dependent on a homogeneous magnetic field, and it results in fewer artifacts than seen with frequency-selective fat suppression techniques.

New developments that provide for long echo-train length, short echo spacing, and half-Fourier computations provide for single-shot imaging that can be performed in patients unable to suspend respiration. However, these methods may result in a decrease in T2 contrast for detecting solid lesions such as HCC.4,5

Breath-hold, T1-weighted images should be performed using a short TR/short TE gradient-echo pulse sequence with a large flip angle. The authors acquire these at two separate echo times for in-phase and out-of-phase imaging (4.4 msec in-phase and 2.2 msec out-of-phase at 1.5 tesla). This improves the detection of either diffuse or focal fatty infiltration and nodular lesions that contain fatty components.

Imaging at short echo times has both benefits and disadvantages. It minimizes the magnetic susceptibility artifacts from transjugular intrahepatic portosystemic shunts (TIPS) and certain embolization coils, which can severely degrade image quality. However, it also decreases the ability to detect susceptibility effects of iron within the liver, whether diffuse or within siderotic nodules. A third gradient-echo pulse sequence with increased echo time and decreased flip angle can be used to increase the magnetic susceptibility effects of iron. This sequence can also be used to determine direction of portal venous flow when saturation bands are placed above and then below the imaging volume.

For MRI of the cirrhotic liver, dynamic, breath-hold, gadolinium-enhanced imaging performed with either two- or three-dimensional gradient-echo pulse sequences is essential.6-8 Contrast-enhanced, frequency-selective fat-suppressed sequences can be used to improve conspicuity of liver lesions, particularly in the setting of fatty infiltration. Fat suppression is also helpful when evaluating the extrahepatic manifestations of cirrhosis, including varices and bowel edema. Other contrast agents are available, including agents that are specifically taken up by hepatocytes or by reticuloendothelial cells. A discussion of these agents is beyond the scope of this article, and the reader is referred to a recent review.9

Contrast-enhanced images should be acquired in at least three phases: the hepatic arterial, portal venous, and equilibrium phases. A successful arterial phase is critical for the detection of small HCC.6,7 Traditionally, hepatic arterial phase images have been obtained following a fixed delay after intravenous bolus of contrast. However, use of a fixed delay often results in a suboptimal hepatic arterial phase in the cirrhotic cohort because of the wide range in circulation times that result from hemodynamic alterations intrinsic to the disease. In these patients, recent developments suggest that using a test bolus to determine an individual's circulation time, or fluoroscopic "real-time" triggering, can help obtain arterial phase imaging more reliably. In addition, an MR-compatible power injector is helpful for injections of the test bolus and main bolus of contrast material, as it provides a precise infusion rate (typically, 2 mL/sec) and dose (0.1 mmol/kg) and allows a technician to perform the entire examination outside of the magnet.

When near-isotropic pixel size (such as lesser than or equal to 2 mm in all three dimensions) can be achieved, 3-D gradient-echo imaging has clear advantages over conventional 2-D. Image sets can be reformatted using multiplanar reconstructions without loss of in-plane resolution. Moreover, angiographic reconstructions such as maximum intensity projections can be obtained for each contrast-enhanced acquisition, resulting in a "free" angiogram and portogram.8

The equilibrium-phase, contrast-enhanced study can be helpful for distinguishing HCC from RN and DN, as some HCC will have a discernable capsule at this stage while RN and DN should not.

RN of cirrhosis have characteristic features on MR imaging that usually allow distinction from HCC, but not always from DN. They invariably have a portal venous blood supply with minimal or no contribution from the hepatic artery. RN are usually isointense with other background nodules on both T1-weighted and T2-weighted images. Less commonly, they may be hyperintense on T1-weighted images and hypointense on T2-weighted images. However, unlike some HCC, RN are almost never hyperintense on T2-weighted images, with the noted exception of those that occur in the setting of chronic Budd-Chiari syndrome10 or those that have undergone infarction.

3D Spin Echo–Type Sequences : 3 Tesla MRI of the Spine

With the higher SNR of 3 Tesla and software developments, 3D spin echo–type sequences with good contrast to noise ratio become possible in a reasonable scan time of 5 to 6 minutes. These MR sequences include sampling perfection with application-optimized contrasts using different flip angle evolutions, CUBE, and volume isotropic turbo spin-echo acquisition. They are somewhat insensitive to magnetization transfer–related cross talk among slices, allow thin-slice multiplanar reconstructions from isotropic voxels, and can produce a variety of contrasts (eg, proton density and T2). They may be used with fat-suppression techniques, such as SPAIR and STIR .

Rodegerdts and associates8 described their experience with 3D sequences of whole-spine MRI at 3 Tesla MRI compared with 1.5 Tesla MRI. They concluded that 3D imaging has a higher SNR on 3 Tesla than on 1.5 Tesla and that better signal can be used to increase the spatial resolution. Of note, the interpatient variability at 3 Tesla was greater than at 1.5 Tesla because of the B1 inhomogeneity related to dielectric effects.

Techniques Used for Spondyloarthropathy : 3 Tesla MRI of the Spine

WBMRI and whole-spine MRI, both somewhat new techniques, are gaining ground in oncology for detection of metastasis, especially in pediatric populations because of their ability to scan the entire spine in only 1 session. They also are being used for spondyloarthropathy (SpA). Multiple phased-array surface coils are connected, negating the need to reposition the patient and making it possible to scan the entire spine in 2 planes in about 20 minutes. WBMRI detects inflammatory lesions outside the sacroiliac joints in a significant number of patients with SpA.


Several other factors support the use of whole-spine MRI in SpA. Whole-spine MRI and conventional MRI show similar results in detecting inflammatory lesions.10,11 Whole-spine MRI is important, because almost 23% of patients with SpA do not have sacroiliitis; therefore, a focused examination of sacroiliac joints could overlook a significant number of pathologies.

In a patient with amyloidosis involving the lumbosacral plexus, coronal maximum intensity projection 3D sagittal sampling perfection with application-optimized contrasts using different flip angle evolutions short tau inversion recovery sequences (A) and axial T2 spectral adiabatic inversion recovery sequences (B) show enlargement of the right L4 and L5 nerve roots (arrows) with amyloid depositions (arrow).
MRN is another application that exploits the higher SNR of 3 Tesla MRI. It consequently gives higher spatial and contrast resolution for evaluating peripheral nerve pathologies, ranging from plexus pathologies to more distal nerve pathologies. Strictly speaking, this technique is separate from spine MRI, but it complements spine imaging, because the cervical plexus and lumbosacral plexus are so closely related to the spine (Figure ).

In patients who have nonlocalizing or nonspecific symptoms and in whom spine imaging is noncontributory, MRN may show nerve abnormality related to injury or inflammation. Defining lesions preoperatively also is useful. In our institution, we use a combination of anatomical (T1-weighted) sequences and 2D and 3D fluid-sensitive sequences (STIR or SPAIR in combination with a 3D turbo spin-echo sequence).

Proton MRS is another technique that benefits from the higher SNR of 3 Tesla MRI. This approach adds metabolic information to the morphological imaging of the bone and soft tissue structures, including the spinal cord. Marliani and colleagues14 demonstrated the feasibility of proton MRS of the cervical spinal cord on 3 Tesla MRI overcoming the challenges from magnetic field heterogeneity around the spinal cord, respiratory and cardiac movement, and the small size of the spinal cord.

Functional (diffusion) imaging has been tested, and good accuracy in differentiating benign from malignant fractures has been reported. In addition, diffusion tensor imaging and tractography is feasible because it provides good fat-suppression techniques, better echo spacing, a higher SNR, and parallel imaging. Although anatomical imaging is useful for detecting macroscopic findings, diffusion tensor imaging provides a way to interrogate tissue microarchitecture and neural integrity.

Safety and MRI

Although MRI is completely safe, it is instructive to consider how the scanner interacts with the patient. To put this section into historical context, in 1980 there were concerns about using field strengths as little as 0.35 T but within 6 years this 'safe' limit had moved up to 2.0 T. Similarly, gradient performances were limited to 3 T/s in the mid-1980s whereas today MRI is routinely performed with gradients exceeding 50 T/s. 
What follows is a summary of each particular safety issue associated with MRI. It is intended to be educational and certainly should not be misconstrued: MRI is entirely safe and I regularly volunteer for scans as part of our research! 

Static Field Effects
The most obvious safety implication is the strength of the magnetic field produced by the scanner. There are three forces associated with exposure to this field: a translational force acting on ferromagnetic objects which are brought close to the scanner (projectile effect), the torque on patient devices/implants, and forces on moving charges within the body, most often observed as a superposition of ECG signal. In the main, sensible safety precautions and the screening of patients means that there are seldom any problems. Of major concern is the re-assessment of medical imaplants and devices deemed safe at 1.5 Tesla which may not have been tested at higher fields. This is becoming an issue as 3.0 T scanners become more commonplace. 
The extension of the magnetic field beyond the scanner is called the fringe field. All modern scanners incorporate additional coil windings which restrict the field outside of the imaging volume. It is mandatory to restrict public access within the 5 Gauss line, the strength at which the magnetic field interfers with pacemakers. 

Gradient Effects
These come under the term 'dB/dt' effects referring to the rate of change in field strength due to gradient switching. The faster the gradients can be turned on and off, the quicker the MR image can be acquired. At 60 T/s peripheral nerve stimulation can occurr, which although harmless may be painful. Cardiac stimulation ocurrs well above this threshold. Manufacturers now employ other methods of increasing imaging speed (so called 'parrallel imaging') in which some gradient encoding is replaced. 

RF Heating Effects
The repetitive use of RF pulses deposits energy which in turn causes heating in the patient. This is expressed in terms of SAR (specific absorption rate in W/kg) and is monitored by the scanner computer. For fields up to 3.0 Tesla, the value of SAR is proportional to the square of the field but at high fields the body becomes increasingly conductive neccessitating the use increased RF power. Minor patient burns have resulted from use of high SAR scans plus some other contributory effect, e.g. adverse patient or coil-lead positioning, but this is still a rare event. 

Noise
The scans themselves can be quite noisey. The Lorentz forces acting on the gradient coils due to current passing through them in the presence of the main field causes them to vibrate. These mechanical vibrations are transmitted through to the patient as acoustic noise. As a consequence patients must wear earplugs or head phones while being scanned. Again, this effect (actually the force on the gradients) increases at higher field and manufactures are using techniques to combat this including lining the scanner bore or attaching the gradient coils to the scan room floor thereby limiting the degree of vibration. 

Claustrophobia
Depending on the mode of entry into the scanner (e.g. head first or feet first) various sites have reported that between 1 % and 10 % of patients experience some degree of claustrophobia which in the extreme cases results in their refusal to proceed with the scan. Fortunately, modern technology means that scanners are becoming wider and shorter drastically reducing this problem for the patient. In addition, bore lighting, ventilation as well as the playing of music all help to reduce this problem to a minimum.

Bioeffects
There are no known or expected harmful effects on humans using field strengths up to 10 Tesla. At 4 Tesla some unpleasant effects have been anedoctally reported including vertigo, flashing lights in the eyes and a metallic taste in the mouth. Currently pregnant women are normally excluded from MRI scans during the first trimester although there is no direct evidence to support this restriction. 
The most invasive MR scans involve the injection of contrast agents (e.g. Gd-DTPA). This is a colourless liquid that is administered i.v. and has an excellent safety record. Minor reactions like warm sensation at the site of injection or back pain are infrequent and more extreme reactions are very rare. 

Nuclear Spin and Behaviour in a Magnetic Field

Electromanetism tells us that a current carrying conductor e.g. a piece of wire, produces a magnetic field encircling it. When the wire is formed into a loop the field acts perpendicular to the surface area of the loop. Analogous to this concept is the field produced by negatively charged electrons orbitting the nucleus in an atom, or the spinning charge of the nucleus itself. This spinning momentum of nuclear charge ('the spin') produces a small magnetic field referred to as a magnetic moment. Under normal circumstances these moments have no fixed orientation so there is no overall magnetic field. However, when nuclei are placed in an external magnetic field, for example a patient placed in the MRI scanner, they begin to align in given directions dictated by the laws of quantum physics. It turns out that in the case of the hydrogen nucleus (a single proton with a spin quantum number, I = ½) that two discrete energy levels (2I +1) are created; a higher energy level where the magnetic moments are opposing the external magnetic field, and a lower energy level in which the nuclei are aligned with the magnetic field. It turns out that a tiny majority of spins are in the latter energy state thereby creating a net magnetisation in the direction of the main magnetic field. The population difference, and therefore the senstivity of the technique, can be altered by reducing the temperature or increasing the field, hence the need for a strong magnetic field, which for modern clinical scanners is between 0.5 and 3.0 Tesla. We refer to this field as B0 to distinguish it from a second field described later on. To put the magnitude of this field into context, 1 Tesla is equal to 10,000 Gauss and the Earth's magnetic field varies from between 0.3 - 0.7 Gauss. 

In terms of classical physics, when the spin is placed in a magnetic field it precesses about that field in a motion analogous to a spinning top. The frequency of precession is governed by the Larmor equation, w0 = gB0. The constant of proportionality in this equation is the magnetogyric ratio with every 'MR visible' nucleus having its own specific value. For the proton, in a field strength of 1.5 T, this frequency is about 63.8 MHz, which is in the radio-frequency (RF) range. 

MRI-guided breast biopsy

During this minimally invasive, image-guided procedure, the patient lays face down on a padded exam table and the affected breasts are positioned into openings in the table. We can adjust the table’s padding for each patient to improve comfort.*

The breasts are then gently compressed between two plates marked with a grid structure, and scanned laterally and medially. This allows us to properly image all quadrants of the breast, the axilla and up into the breast wall.

Using targeting software, the radiologist measures and pinpoints the position of the breast tumor with respect to the grid. This helps us calculate the position and depth of the needle placement for the biopsy.

In preparation for the biopsy, we first inject a local anesthetic into the breast to numb it. Next, the radiologist inserts a core needle and advances it to the location of the tumor. We then use magnetic resonance (MR) imaging to verify the tumor’s position.

Once the tumor’s position has been confirmed, a vacuum-assisted needle uses vacuum pressure to pull tissue from the breast into the sampling chamber. The tissue samples are then taken to the laboratory for pathology testing.

Advantages of MRI-guided breast biopsy

The MRI device and targeting software allows our radiologists to position the coils as close to the breast as possible to achieve higher quality images and faster scan times.

The ergonomic design and visco-elastic padding of the table provide greater comfort throughout the procedure.*

The tabletop design allows the patient to position her arms at her side instead of over the head, resulting in less extension of the breast muscle and increasing the amount of breast wall in the imaging cone.*

During the biopsy, the needle rotates positions without having to withdraw and reinsert it, allowing us to collect additional tissue samples (generally eight to 10 samples).

Steady-state versus spoiled gradient echo imaging in Bright Blood Cardiac MRI Sequences

In gradient echo (GRE) imaging, the TR is often shorter than the T2 of most tissues, and the transverse magnetization will not have fully decayed before the next RF pulse. Thus, there will be residual transverse magnetization that adds T2 contrast (in addition to T1 contrast) to the image. This additional T2 contrast is undesirable for many applications, as the T1 and T2 contrast may be competitive. For example, a liver lesion that is hypointense on T1 and hyperintense on T2 may be isointense with both T1 and T2 weighting. To achieve T1 weighting with a short TR GRE sequence, spoiling the residual transverse magnetization is necessary. This spoiling can be accomplished with an RF pulse or gradients. The majority of fast GRE sequences used in noncardiac clinical MRI are spoiled.

In steady-state GRE sequences, spoiling is not performed, and residual transverse magnetization is retained. The retained residual transverse magnetization increases the signal-to-noise ratio (SNR) of steady-state sequences relative to spoiled sequences. The image contrast will depend on the T2-to-T1 ratio. As stated previously, this is undesirable for many applications. In steady-state sequences, only fluid and fat will have a high signal (fluid and fat have comparable T1 and T2 times, while in most other tissues, T2 time is much shorter than T1 time). However, in bright blood cardiac MRI, hyperintense blood relative to other tissues is exactly what is needed; thus, steady-state GRE sequences are optimal for cine cardiac imaging (cMRI).

The sequences used in cardiac imaging are balanced SSFP sequences. Different trade names for these sequences are TrueFISP (Siemens), FIESTA (GE), and balanced FFE (Phillips). These sequences are very fast and have a high SNR, but the T2-to-T1 image contrast limits the role of these sequences for noncardiac applications.

SSFP cine MRI has largely replaced spoiled GRE cine MRI for evaluation of cardiac function. SSFP sequences do not depend on flow; they have a higher SNR; and they are faster. Spoiled GRE sequences are T1 weighted and depend on through plane flow enhancement (similar to time-of-flight MR angiography) to generate contrast. The blood may become saturated if the flow is slow or the TR is short. Thus, spoiled GRE cine MRI does not allow for the use of very low TRs, because there is not enough time for saturated blood to be replaced by unsaturated blood between excitation pulses.

With SSFP sequences, blood signal is dependent on intrinsic contrast rather than inflow effects, and TR can be as short as possible. SSFP cine MRI can be almost 3 times as fast as spoiled GRE cine MRI. In addition, the SSFP sequence has a higher SNR due to the residual transverse magnetization. This is particularly true at low TRs. With spoiled GRE sequences, SNR decreases with decreasing TR. With SSFP sequences, SNR is high even at low TRs, because residual transverse magnetization increases with shorter TRs.