A Numerical Systematic Review and Meta-Analysis of Diagnosing the Vibration Modes of the Cylindrical Shell in the MRI Machine

The Magnetic Resonance Imaging (MRI) machine is a subset of nuclear magnetic resonance imaging technology that produces images of the body using magnetic field gradients. The MRI Machine has two components: the computer-based control centre room and the adjacent MRI machine room where the patient undergoes the scan. This study aimed to assess the quality and reliability of YouTube videos about MRI machines, MRI scans, and MRI claustrophobia and compare the quality and reliability of the videos among different types of uploaders. Methodology: The YouTube Search Algorithm and a Google Sheets questionnaire were used to evaluate 10 videos that satisfied the inclusion criteria of the study. The video analytics included were title, number of views, likes and dislikes, comments, duration, source, and content. The quality of each video was established using the Global Quality Score (GQS), Reliability Score, and Video Power Index (VPI), where each quantifier went through statistical analysis using SPSS software, version 21.0 (IBM Corp., Armonk, NY) to determine if there was any significance. In order to determine statistical differences between the groups, the Kruskal-Wallis test was used on the quantifiers GQS, reliability score, and VPI to generate p-values. The p-valuefor VPI is 0.467, GQS is 0.277, and reliability is 0.316. All the p-values are greater than 0.05, showing that there is no statistical support for any significant difference between the groups in their VPI, GQS and reliability scores. YouTube videos with high-quality and reliable information on MRI machines, MRI procedures, and claustrophobia, especially those uploaded by clinicians and hospitals, can provide correct information, helping patients decide to undergo these procedures and alleviate claustrophobia.

Magnetic resonance imaging (MRI) gradient coils produce acoustic noise due to coil conductor vibrations caused by large Lorentz forces. Accurate sound pressure levels and modeling of heating are essential for the assessment of gradient coil safety. This work reviews the state-of-the-art numerical methods used in accurate gradient coil modeling and prediction of sound pressure levels (SPLs) and temperature rise. We review several approaches proposed for noise level reduction of high-performance gradient coils, with a maximum noise reduction of 20 decibels (dB) demonstrated. An efficient gradient cooling technique is also presented.

Paul C Lauterbur

Electro-magnetic fields and wireless technology are part of modern life. The use of Magnetic Resonance Imaging (MRI) machines for clarification of internal human structures and function in healthcare is increasing. The rapid development of wireless devices, their miniaturization and their application as clinical tools creates an expanding intersection zone. Although safety standards for devices in MRI machines have been previously published, it is not clear that newer wireless technologies, including devices used in Medical Body Area Networks (MBAN) have been rigorously tested or disclosed. We undertook a review of the clinical scientific literature and the United States Food and Drug Administration adverse events database to discover whether this is a significant issue. There are currently no published studies specifically addressing the safety of wireless devices potentially used in MBAN in MRI machines. We suggest the addition of a research track to clarify the safety of MBAN devices in MRI machines. Informed design of current and future MBAN components, devices and systems can avoid potential patient adverse events due to the un-intended consequences of the concurrent use of these technologies in MRI machines.

To examine peripheral nerve stimulation (PNS) thresholds for normal human subjects in magnetic resonance imaging (MRI) gradient coils, and determine if observed thresholds could be predicted based on gross physiologic measurements. PNS thresholds for 21 healthy normal subjects were measured using a whole-body gradient coil. Subjects were exposed to a trapezoidal echo-planar imaging (EPI) gradient waveform and the total change in gradient strength (DeltaG) required to cause PNS as a function of the duration of the gradient switching time (tau) were measured. Correlation coefficients and corresponding P values were calculated for the PNS threshold measurements against simple physiologic measurements taken of the subjects, including weight, height, girth, and average body fat percentage, in order to determine if there were any easily observable dependencies. No convincing correlations between threshold parameters and gross physiologic measurements were observed. These results suggest it is unlikely that a simple physiologic measurement of subject anatomy can be used to guide the operation of MRI scanners in a subject-specific manner in order to increase gradient system performance while avoiding PNS.

DUCAS is a computer code using an arbitrary magnetic field distribution as input and producing a sheet current distribution as output. The code was improved for application to the magnetic resonance imaging (MRI) gradient coil (GC) design. DUCAS computes the current potential (CP) distribution, and the coil conductors are placed along the contours, which are current flow lines. An improved algorithm with weights on node CPs and input initial CPs is described for use with DUCAS. The algorithm obtains the CP distribution by summing singular value decomposition (SVD) eigendistributions with the initial CP distribution. The weights modify the eigendistributions. The initial CP distribution is used to get smooth flow lines and then coil patterns. The improved DUCAS is suitable for avoiding a complex current distribution and to obtain coil patterns that are easy to manufacture. The initial CP distribution is assumed first, then DUCAS compensates the CPs for the given magnetic field distribution with the weights. This paper describes the computational algorithm and shows a GC design obtained with it.

To check the bioeffects of the components of magnetic resonance imaging (MRI). MRI is based on an assumed harmless interaction between certain nuclei in the body when placed in a strong magnetic field and radio wave fields. There are three key factors actuating on the examining body: a powerful static magnetic field (SMF), magnetic gradient fields (MGFs), and pulsed radiofrequency (RF) radiation. In vitro cells (L-132 cells) were used as biosensors, and different cellular compounds were used as biomarkers (heat shock proteins [HSPs] and their messenger ribonucleic acids [mRNAs], calcium, and adenosine-3',5'-cyclic monophosphate [cAMP]). The biosensors were placed in the bore of a 1.5-T MRI machine and the different electromagnetic fields were operated. HSPs and their mRNAs and cAMP did not respond to SMF, MGFs, or RF radiation. RF radiation increased cytosolic calcium concentration (18%, P < 0.05). Although MRI procedures do not induce any cellular stress response, it may cause an unfathomable calcium increase in vitro. Although the in vitro experimental conditions are not totally comparable to clinical situations, the usefulness of the in vivo biological dosimetry, circulating leukocytes as biosensors, and HSPs and/or calcium as biomarkers is suggested.

Background: This research on the performance testing of the Magnetic Resonance Imaging (MRI) machine was conducted using performance tests on three different parameters at Kasih Ibu Hospital Saba. This research aims to determine the performance of the MRI machine at Kasih Ibu Hospital Saba in Bali Province and establish baseline data for future testing. Method: This type of research is quantitative and uses a survey approach. The tools and materials used in this study are a 0.35 Tesla MRI machine, an ACR phantom, and a head coil. Data were collected by conducting nine MRI performance testing procedures using the ACR (2015) guidelines and then analyzed using international standards issued by ACR (2015). Result: In the visual checklist testing, all observed parameters were over 30 days and functioned well, For the Artifact Evaluation test, the author evaluated all scanned images and found no artifacts. The author analyzed the measurements from the first slice of T1W and T2W images for the Slice Thickness Accuracy test. The measurement results for the T1W image were 4.6mm, and for the T2W image, they were 5.57mm. The Slice Thickness Accuracy testing standard set by ACR in 2015 is 5mm ± 0.7mm. Therefore, based on the measurement results from the T1W and T2W images, the MRI machine at Kasih Ibu Saba Hospital meets the established international standards. Conclusion: Based on the results of the visual checklist test, the following results were obtained: all parameters on the visual checklist functioned well during the 30-day observation. Meanwhile, the two tests conducted with the phantom on the MRI machine at Kasih Ibu Hospital Saba the following results are: the artifact evaluation test and the slice thickness accuracy test met the international standards set by ACR in 2015.

There is higher disease incidence and worse outcomes in rural America when compared to urban America. In states like Wisconsin, where 32.9% of the population resides in rural areas, this is particularly worrisome. The Center for Healthcare Quality and Payment Reform found that 30% of rural hospitals in the US are at risk of closing due to financial instability. A substantial cost to rural hospitals is the provision of radiologic services. Thus, the study investigated if a disparity exists in availability of magnetic resonance imaging (MRI) and computed tomography (CT) machines among Wisconsin's urban and rural county hospitals. Wisconsin hospitals were asked how many MRI and CT machines were carried at their facility. This information was compiled in a spreadsheet and cross-referenced with the county in which it resided, along with the county's population, urban-rural classification, and land area in square miles. We found that the state of Wisconsin compared favorably with the national average in terms of the number of persons and square miles per MRI and CT machine. When comparing Wisconsin counties based on their urban-rural classification, a disparity exists in rural counties regarding square mileage per CT and MRI machine. With distance for service creating a barrier to accessibility, rural county residents would benefit from more in-hospital MRI and CT machines. Based on these findings, further research is warranted to investigate the potential vulnerability of other rural populations regarding accessibility to radiologic resources.

Magnetic resonance imaging (MRI) devices provide magnetic field gradients, which can be used to remotely actuate magnetic robots that may one day carry out medical tasks, such as diagnosis, drug delivery, or therapeutic-laser-based procedures, in addition to the high-resolution tissue images for diagnosis. However, in comparison with magnetic systems that are custom designed solely for magnetic actuation, the magnetic environment of an MRI device is constrained by the requirements for imaging, which reduces the number of active degrees of freedom available for magnetic actuation. Moreover, the current MRI-powered untethered robots are limited to translational magnetic pulling in three dimensions only. In this article, we propose a design for an untethered magnetic robot that can rotate in a three-dimensional liquid volume in magnetic environments, like those in MRI devices. We demonstrate rotational actuation of our robot inside a commercially available MRI gradient coil. Two kinds of near-neutrally buoyant robot designs are proposed, where each design has particular advantages. Design methodologies, analysis of rotational performance, closed-loop orientation control of up to 2.51 rad of orientation change with maximum net displacement of 18.4 mm, and angular velocity control (0.2-0.63 rad/s) of these robot designs are presented.

Gradient coil design and acoustic noise control in magnetic resonance imaging systems

Magnetic resonance imaging (MRI) is based on nuclear magnetic resonance, which was independently discovered in 1946 by Bloch et al. (1946) and Purcell et al. (1946). Therefore, the first scanners to use nuclear magnetic resonance to obtain medical images were called nuclear magnetic resonance scanners. The phenomenon of nuclear magnetic resonance is based on the interaction between an external magnetic field and nuclei. However, patients were concerned about the use of the word “nuclear” because they associated “nuclear” with radioactivity, even though no radioactivity or ionizing radiation is involved; therefore, the term “MRI” came into use. MRI shows contrast between soft tissues and has high spatial resolution (Kuperman, 2000). This technique has rapidly evolved over the last 30 years to become an important imaging modality. Approximately 60 million MRI scans are carried out throughout the world every year. Magnetic resonance occurs when a magnet interacts with a magnetic field (Oldendorf and Oldendorf, 1988). MRI is possible because tissue in the body is composed of lots of water. Water molecules are composed of hydrogen nuclei (protons). A hydrogen atom comprises a nucleus with one proton and one electron that orbit the nucleus. The proton is able to spin and behaves like a small magnet. Hydrogen has a highly magnetic nucleus and makes up two thirds of the atoms in tissues. Hydrogen nuclei are the best target for in vivo MRI because, among all types of nuclei in tissues, they provide the highest nuclear magnetic resonance signal and achieve good contrast between different tissues in vivo. (Kuperman, 2000). In the nucleus of hydrogen, movement in electric charge results in a magnetic field. The hydrogen nucleus spins about an axis, and therefore, this moving charge behaves similar to current in a loop of wire, producing a magnetic field. This is termed nuclear magnetism, which can be used for imaging. When the hydrogen nucleus is exposed to a strong magnetic field, it has properties that are comparable with those of a compass needle in the Earth's magnetic field. First, the hydrogen nucleus tends to align with a magnetic field. Second, the hydrogen nucleus has a resonant frequency proportional to the external field strength. Third, the hydrogen nucleus absorbs energy if this energy is at the resonant frequency, and this energy will subsequently be re-emitted. In the presence of a gradient magnetic field, the location of hydrogen nuclei in the gradient is indicated by their resonant frequency. When an atomic nucleus becomes a resonating magnetic object, this is termed nuclear magnetic resonance. When the nucleus in a hydrogen atom is stimulated by radio waves, which have a weak magnetic field, the absorption of energy from radio waves by a magnetic atomic nucleus constitutes the phenomenon of nuclear magnetic resonance. The term “resonant frequency” is the number of times per second that the hydrogen nucleus wobbles or “precesses” (oscillation in a wobbling motion). As mentioned above, some protons will align in the direction of the magnetic field, but some will align in the opposite direction to the magnetic field. Magnetic fields from many of these protons will be cancelled out, but more protons will tend to be aligned with the main magnetic field, which produces a “net magnetization” that is aligned parallel to the main magnetic field. This net magnetization is the source of the MR signal (Pooley, 2005). When energy is absorbed from a radiofrequency pulse, this net magnetization rotates away from the longitudinal direction. The duration and strength of the radiofrequency pulse determines the amount of this rotation, which is called the flip angle. When net magnetization is rotated into the transverse plane (transverse magnetization) by a radiofrequency pulse, this is called a 90° radiofrequency pulse. When net magnetization is rotated 180° into the longitudinal direction (longitudinal magnetization), this is called a 180° radiofrequency pulse. The main components of an MRI system comprise the following: (1) a strong magnet for generating a static magnetic field; (2) a gradient system comprising three coils for producing a linear field in the X-, Y-, and Z-directions (see “Basic Scans” below for discussion of these axes) and the corresponding amplifiers; (3) a radiofrequency transmitter including a transmit coil in the scanner; (4) a sensitive radiofrequency transmitter receiver for picking up and amplifying the MR signal; (5) coils for either receiving or transmitting; and (6) computers for regulating the scanner and the gradients, for creating the MR images (array processor), and for coordination of all processes. (Weishaupt et al., 2006). The magnetic field produced by the magnet must have adequate strength, which usually ranges from 0.1 to 3 Teslas (T) in medical imaging. Most of the MRI scanners used operate at 1.5 T, but 3-T scanners are becoming more common in the clinical setting. In addition, some 7-T scanners are being used in universities, particularly for small animal imaging research. The MRI scanner contains a radio transmitter for producing stimulation. This scanner produces the strong fields required to align the hydrogen nuclei. The scanner also produces gradient fields to localize the nuclei. An individual nucleus emits a signal that is too weak for measurement by an MRI scanner, but the radio waves re-emitted by millions of nuclei can be sensed by an antenna of the scanner after stimulation. Magnetization needs to be large enough to acquire a signal because the MR signal is weak (Weishaupt et al., 2006). An imaginary construct called the magnetization vector is used for describing the behavior of hydrogen nuclei during MRI. This is a single vector, representing the behavior of all of the hydrogen nuclei in a small region of tissue. The magnetization vector uses a coordinate system in which the X-, Y-, and Z-axes correspond to those of the magnetic field of the scanner. Imaging a transverse slice of tissue is the most common in MRI (Oldendorf and Oldendorf, 1988). To image a transverse slice, the slice must first be isolated from other tissues. This is achieved by the scanner creating a magnetic gradient longitudinally through the body, along what is called the Z-axis. The direction parallel to the main magnetic field is defined as the longitudinal direction. The hydrogen nuclei become spatially encoded when this gradient is present. In adjacent tissue slices, hydrogen nuclei are unstimulated. A second gradient is then produced at right angles to the Z-axis after the first Z-axis gradient is turned off. This results in fields of different strengths. Consequently, the radio signal emitted from the plane comprises a mixture of frequencies. The amount of hydrogen along each line in the tissue slice can be obtained by measuring the radio signal strength at different frequencies. To obtain sufficient data for forming an image, a tissue slice needs to be examined multiple times. Therefore, following a stimulating pulse, the Z-axis gradient is switched off again, and a transverse gradient is given across the tissue slice, but in a slightly different direction to the previous gradient. The plane perpendicular to the Z direction is called the transverse plane (X–Y plane). This process must be repeated so that the single slice of tissue is finally cut into several hundred sets of parallel lines, and each set crosses at a different angle on the plane of the slice. The first gradient can also be transversely applied across the body for producing a sagittal section or it can be applied front-to-back to obtain a coronal section. The final step of the imaging process is achieved by computer. To obtain an image, the Fourier transform is performed. Hydrogen nuclei are affected by their chemical environment. This environment alters the behavior of hydrogen nuclei, leading to changes in the radio signal quality emitted by tissues. When the hydrogen nucleus is exposed to a magnetic field, it is subject to small local variation in the strength of the magnetic field. MRI has good diagnostic potential because of its ability to determine differences in the magnetic environment surrounding hydrogen nuclei. Changes in the magnetic environment alter the signal received after excitation. Two types of behavior of the hydrogen nucleus are influenced by the local magnetic environment. These are called time constants T1 and T2, and are also known as relaxation times because they represent the time it takes for the emitted signal to fade following stimulation. Longitudinal magnetization is zero following a 90° radiofrequency pulse. Longitudinal magnetization is then slowly restored, and this process is known as longitudinal relaxation or T1 relaxation (Weishaupt et al., 2006). More precisely, T1 is defined as the time that it takes for longitudinal magnetization to reach 63% of its final value when there is a 90° radiofrequency pulse. The rate at which T1 relaxation occurs is different for protons in different tissues. This situation results in contrast in T1-weighted images. Fluids have a long T1 (1,500–2,000 ms), water-based tissues have an intermediate T1 (400–1,200 ms), and fat-based tissues have a short T1 (100–150 ms). An example of this type of contrast can be found in the brain. White matter has a short T1 time and rapidly relaxes. In contrast, cerebrospinal fluid has a long T1 and slowly relaxes. The T1 time in gray matter is intermediate and has an intermediate relaxation rate. If an image is produced when these relaxation curves are widely separated, this results in an image with high contrast between the tissues. Consequently, white matter is shown as light pixels, cerebrospinal fluid shows dark pixels, and gray matter shows pixels with shades of gray. A representative T1 image is shown in Fig. 1. T2 represents how long transverse magnetization would last in a perfectly uniform external magnetic field. In other words, T2 is a measure of the length of time that resonating protons are coherent or rotate in phase after a 90° radiofrequency pulse. This process is known as transverse relaxation or T2 relaxation. More precisely, T2 is defined as the time that it takes for transverse magnetization to decay to 37% of its original value. Transverse relaxation is defined as the decay in transverse magnetization because spins lose coherence (known as dephasing). Transverse relaxation is different from longitudinal relaxation in that the spins exchange energy with each other instead of energy being dissipated to their surroundings. Similar to T1, different tissues have different T2 values and dephase at different rates. Fat is differentiated from water, but in contrast to T1, in T2, fat is darker, and fluid is lighter. Representative T2 images are shown in Figs. 2 and 3. The features that affect signal intensity or brightness of a biological tissue on an MR image, and consequently image contrast, are proton density and T1 and T2 times (Weishaupt et al., 2006). These features can greatly vary between tissues. By varying which of these parameters is emphasized in an MR sequence, this determines how the resulting images are different in their contrast. This process is the basis for the fine discernment of soft tissue and diagnostic potential of MRI. T1 and T2 relaxation are independent of each other, but occur simultaneously. In addition, the time to echo and repetition time (see “Types of Pulse Sequences” below) are useful for controlling the amount of “weighting” of T1 and T2 effects in the image (Pooley, 2005). MRI using a single pulse shows the distribution of hydrogen throughout the body and is not useful for diagnostics (Oldendorf and Oldendorf, 1988). However, MRI is usually used to address more complex issues. To investigate these complex issues, pulse sequencing is carried out instead of a single pulse. For pulse sequencing, two or more radio pulses are applied to tissue in quick succession. A parameter involved in pulse sequencing is called the repetition time, which is the time taken to go once through the pulse sequence. Researchers need a lot of experience to interpret scans made with various pulse sequences. This is because many different images can be obtained of the same scan slice with different pulse sequences. Many pulse sequences have been developed, but only some of the more common ones used in MRI are discussed below. Future development of MRI is likely to involve new pulse sequences because many variables can be used. The spin echo pulse sequence is a common type, in which a 90° pulse is followed by a 180° pulse. Following the 90° radio frequency pulse, protons in phase begin to diphase. The protons begin to rephase after the 180° pulse. This rephasing of the spins causes formation of an “echo”, termed a spin echo. The time between the peak of the 90° radio frequency pulse and the peak of the echo is termed the time to echo. For clinical MRI, scans are usually created from spin-echo pulse sequences. Another type of spin echo pulse sequence is multiecho spin echo, in which multiple 180° radio frequency pulses are performed to form multiple echoes. Each of these echoes is used for constructing a separate image data set with different contrast weighting. Turbo-spin echo, otherwise known as fast-spin echo, also uses multiple 180° radio frequency pulses to create multiple echoes. However, rather than each echo producing a different image data set, all of the echoes are included to achieve a single image data set in a faster time. The turbo-spin echo pulse sequence is useful for producing T1 and T2 contrast weighting (Figs. 1-3). Each echo still occurs at a different time to echo and thus has a different contrast weighting. Another common pulse sequence used clinically is inversion-recovery. This type of sequence measures T1 and mostly eliminates T2 effects from the scan. This enables greater anatomical details and better gray-white matter contrast than T2-weighted scans. Inversion-recovery is used to suppress undesired signals in MRI (e.g., signals from fat or fluid). Selecting time to echo and repetition time can still be performed to control contrast weighting. The inversion-recovery pulse sequence is different from the spin-echo pulse sequence in that a 180° radio frequency pulse is given before the regular spin-echo pulse sequence of 90°. There are many specialized MRI techniques, such as diffusion-weighted MRI, fluid attenuated inversion recovery (FLAIR) (Fig. 4), time-of-flight magnetic resonance angiogram (Fig. 5) magnetic resonance spectroscopy, functional MRI, magnetic resonance angiography, T1rho MRI, real-time MRI, and magnetization transfer MRI. However, discussion of these techniques is outside the scope of this article. Some of these techniques will be covered in future Insight articles. Contrast agents enhance contrast in MR images by altering T1 and T2 relaxation times (Kuperman, 2000). They are often administered in MRI for improving assessment of local physiological and anatomical conditions or for improving detection of malignancy. Contrast agents are usually administered internally, and should have low toxicity with easy excretion from the body. Contrast agents for MRI indirectly affect the signal by interacting with hydrogen nuclei. Many MRI contrast agents change the T1 and T2 relaxation times in tissue via dipole-dipole interaction with water protons. Some of the more commonly used contrast agents are discussed below. Paramagnetic ions with unpaired electrons (e.g., Gd 3+ and Mn 2+) are often applied as MRI contrast agents because of their high relaxivities. However, these agents can be toxic, and to reduce this effect for in vivo use, they can be chelated to particular molecules. Gadolinium-diethylenetriaminepentaacetic acid (DTPA) is the most well-known MRI contrast agent (Fig. 6). Gadolinium-enhanced MRI is useful for detecting tissue abnormalities (e.g., brain and breast). Gadolinium-DTPA is a chelate composed of gadolinium ions and DTPA. Gadolinium-DTPA has a high margin of safety (Kanal et al., 1990), with a low prevalence of adverse reactions (2.4%). There are two types of iron oxide contrast agents, including superparamagnetic iron oxide and ultrasmall superparamagnetic iron oxide. These contrast agents comprise suspended colloids of iron oxide nanoparticles. When they are injected while imaging, they reduce the T2 signals of absorbing tissues. Superparamagnetic iron oxide is used for clinical liver imaging. Superparamagnetic iron oxide is detectable at much lower concentrations than with paramagnetic agents (e.g., gadolinium). Superparamagnetic iron oxide is nontoxic to cells, is biodegradable in vivo (Arbab et al., 2005; Farrell et al., 2009), and has been shown to be safe for intravenous administration in humans (Richards et al., 2012). The desire to improve enhancement of lymph nodes on MRI led to the development of ultrasmall superparamagnetic iron oxide particles from size fractionation of polydispersed superparamagnetic iron oxide (Bernd et al., 2009). MRI is applied in imaging every area of the body. In particular, MRI is useful for tissues with many hydrogen nuclei and little contrast density (e.g., brain and muscle). Clinically, MRI is used to discriminate pathological tissue from healthy tissue. T2-weighted sequences are especially sensitive for pathology, and are useful for distinguishing pathological tissue from normal tissue. MRI has a number of advantages as follows (Oldendorf and Oldendorf, 1988). (1) Tissue can be characterized in various ways using MRI. (2) MRI can image any plane, as discussed above. (3) The radio signals and magnetic fields used in MRI do not interact with bone. Compact bone is mostly not able to be seen with MRI because it has a low water content. Only fatty bone marrow is able to be imaged, being the only bone structure that can be visualized. (4) Contrast medium is not always needed. (5) Finally, unlike computed tomography scans and X-rays, MRI does not use ionizing radiation. MRI has comparable resolution to computed tomography but has much better contrast resolution. There are some disadvantages of MRI as follows. (1) MRI is complex and expensive. (2) MRI involves long scan times. (3) Noise, isolation, and confinement can be stressful to patients while scanning. (4) Patients who have pacemakers and metallic artifacts are unable to undergo MRI. There are various health and safety concerns for both patients and doctors/researchers because of the way in which clinical MR imaging is currently performed. These concerns can be categorized into the following four main areas: (1) static magnetic field effects (e.g., biological effects and safety considerations [physical effects] related to attractive forces of the static magnetic field; (2) effects from time-varying extremely low frequency magnetic fields (e.g., altered synthesis of DNA); (3) effects of rapid changes in magnetic field gradients, which can induce voltages in the body; and (4) other concerns that are not directly related to magnetic fields (e.g., contrast agents and psychological effects; see http://www.mrisafety.com/safety_article.asp?subject=172) (Kanal et al., 1990). Safety guidelines for occupational exposure to MRI have been published in the USA (e.g., ACR Guidance Document for Safe MR Practices, Sentinel Event Alert #38 by the Joint Commission, and MRI Design Guide by the United States Veterans Administration) and Europe (e.g., EU directive 2004/40/EC on physical agents [electromagnetic fields]). Exposure limits are usually expressed as root mean square values. In most guidelines, restrictions are established to avoid short-term adverse effects and are defined in terms of root mean square-induced electric field components in tissue. However, despite guidelines being issued, little is known regarding the actual levels of electromagnetic field exposure in workers involved with using MRI (McRobbie, 2012). The author thanks Angela Harrison and Anna-Maria Lydon at the Centre for Advanced MRI, the University of Auckland, and Ayton Hope for kindly providing the images.

Brain imaging has the potential to advance our understanding of human health and to improve diagnosis and treatment of neurological diseases. Inspired by key questions in neuroscience and medicine, it becomes extremely important to develop statistical methods that can accurately and efficiently recover useful quantitative information from large amounts of brain images. The underlying computational issues are challenging and often hampered by uncertainties in imaging acquisition parameters, the variability of human anatomy and physiology, as well as the nature of the imaging data to be handled such as the presence of noise and correlation, and the sample and data sizes, and so on. Structural and Functional MRI (sMRI and fMRI) Among the varieties of brain imaging modalities, magnetic resonance imaging (MRI) is primarily a noninvasive imaging technique used in radiology to visualize the brain’s structure and function. Two main forms of MRI include: Structural MRI (sMRI) images the anatomy and strucure of the brain (Symms et al., 2004) and provides detailed pictures of the brain’s size and shape; functional MRI (fMRI) identifies active regions, patterns of functional connectivities during either tasks specifically designed to study various aspects of brain fundtion or during the resting state (Martijn et al., 2010). The MRI machine is, in essence, a big magnet. As the subject lies in its magnetic field, invisible radio waves are released around the subject. This will result in harmless radio waves bouncing off the different substances that make up the brain. The radio waves are then detected by a computer, which transforms the data into images of the brain’s structure and activity. In fMRI, as the subject lies in the MRI machine, simple tasks are given; the MRI then maps what parts of the brain are most active during those tasks compared with activity while the brain is at rest. This allows researchers to understand how the brain functions. This information is used together with the data from the sMRI data to reveal a comprehensive picture of brain structure and function that fit in the overall studies or to allow us to understand how the healthy brain works. The informaiton and fusion of structural and functional MRI can also improve our understanding and the treatment of neurodegenerative diseases and mental disorders such as Alzheimer’s disease and schizophrenia. Brain Morphometry Analysis with Hypothesis Testing from Structural MRI Structural MRI (sMRI), or simply called MRI, scans are usually stored in the format of threedimensional (3D) voxels. There are several procedures for MRI post-processing, and the two

For many years, people working near strong static magnetic fields of magnetic resonance imaging (MRI) machines have reported dizziness and sensations of vertigo. The discovery a decade ago that a sustained nystagmus can be observed in all humans with an intact labyrinth inside MRI machines led to a possible mechanism: a Lorentz force occurring in the labyrinth from the interactions of normal inner ear ionic currents and the strong static magnetic fields of the MRI machine. Inside an MRI, the Lorentz force acts to induce a constant deflection of the semicircular canal cupula of the superior and lateral semicircular canals. This inner ear stimulation creates a sensation of rotation, and a constant horizontal/torsional nystagmus that can only be observed when visual fixation is removed. Over time, the brain adapts to both the perception of rotation and the nystagmus, with the perception usually diminishing over a few minutes, and the nystagmus persisting at a reduced level for hours. This observation has led to discoveries about how the central vestibular mechanisms adapt to a constant vestibular asymmetry and is a useful model of set-point adaptation or how homeostasis is maintained in response to changes in the internal milieu or the external environment. We review what is known about the effects of stimulation of the vestibular system with high-strength magnetic fields and how the understanding of the mechanism has been refined since it was first proposed. We suggest future ways that magnetic vestibular stimulation might be used to understand vestibular disease and how it might be treated.

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