Some fonts have serifs, the slight projections which finish off a stroke of a letter. A common serif font is Times New Roman. Fonts without serifs are called sanserif or simply sans (for example, Arial). The differences between the lettershapes of sanserif typefaces is reduced, making them legible but not very easy to read in larger blocks of text. Adjustments are sometimes made to increase readability. So-called humanist sans fonts retain the clean lines of sans fonts without losing readability or familiarity, by retaining a number of features similar to letters written by the human hand: axis, aperture, modulation, and so on. Gill Sans is such a font.

Philips say of their typography “our typography is inviting and highly legible and has enduring style”. A recent example:

The typeface was designed in 1931 by Eric Gill, and was inspired by a font designed by his teacher, Edward Johnston, for the London Underground. Gill wanted to improve upon Johnston Sans, remarking that ‘some of these letters are not entirely satisfactory, especially when it is remembered that, for such a purpose, an alphabet should be as near as possible “fool-proof”… as the philosophers would say – nothing should be left to the imagination of the sign-writer or enamel-plate maker.’ Whether this motivation behind the design of Gill Sans is relevant today in the context of computer-aided text-generation, is open to argument.

Modern alternative fonts following in the tradition of Gill Sans exist, such as Bliss and FB Agenda. Nevertheless, Gill Sans remains very popular, and has been called the ‘Helvetica of the UK’, meaning it is seen everywhere (as Helvetica is, in US cities). It can be seen used in many companies’ corporate identities and products, including: the BBC, Network Rail, the Church of England, and the British Government which formally adopted Gill Sans as its standard font in 2003. Outside of the UK it is used extensively too: by Monotype Imaging, United Colors of Benetton, Saab, even the G20 group of economies. Its popularity speaks of its flexibility: it has an efficient, professional look but remains comfortable to read.

It has been said that setting body text in Gill Sans requires a sure sense of color [the overall darkness of the type set in mass, which not the same as the weight of the face] and measure [overall width of a textblock]. There are some residual serifs in the lowercase in some of the weights (a, g), and Philips use a version of the font with an alternative, serifed, figure 1. The x-height is lower than competing contemporaries (e.g. Futura), which would reduce readability were it not for the humanist aspects of the font design: variable aperture (compare the openings on the c and e glyphs); and slight contrast in the face [thick vs. thin strokes of letters] which also adds its warmth, as opposed to sans fonts with no stroke modulation.* The significant differences between Gill Sans weights [light, regular, italic, bold, ultra bold, condensed etc] lends the font to many different publication styles and contexts.

* The slight contrast in the regular Gill Sans face does not permit attribution of a clear axis, and so in this aspect it does not adhere to all the characteristics of a humanist sans.

A recent example:

The Siemens font family was designed in 2008 by Hans-Jürg Hunziker, and is delivered by the URW++ type foundry.  The font comes in three typefaces (serif, sans and slab), with eighteen styles in total including all the weights and styles. They were conceived ‘as a set of related modern, mechanistic and lineal faces’, echoing Siemens’ relationship with technology. Siemens intend (pdf) the font to be ‘a distinct typographic style that reflects our character and resonates with our values and beliefs as a company’, and to be ‘exceptionally legible’.

The serif font, seen above, has a distinct style suitable for titles and headings, and the slab and sans versions are complimentary, maintaining the same structure and so adding harmony to the typeface family. The slab face has the same stroke weight as the serif. The sanserif glyphs have reduced contrast compared to the serif, adding to readability when set in body text.

The serif font has: modulated stroke; rationalist [vertical] axis; abrupt trapezoidal serifs (on the serif and slab typefaces); large aperture; flat stroke ends and sharply-modeled serif-like terminals (cf. the endpoints of s and f).

Strokes are cut on the diagonal which helps to prevent the the face appearing too impersonal. The gentle modeling and blunt serifs give the font a good chance of surviving the indignities of low resolution. Subtle modulation of the downstrokes on the serif and sans typefaces give the font added personality at higher resolutions, balancing its industrial design.

A recent example:

GE Inspira was originally designed by Mike Abbink who built on early concepts by Patrick Giasson, in 2005. Inspira comes in four fonts, with various purposes [general use (Inspira Regular); small text (Inspira Book); screen presentation (Inspira Pitch, i.e. bold); and small caps]. The font is a very prominent part of GE’s brand expression, claiming to be ‘derived from the curves and the classic hand drawn character of the monogram’ [logo]. The intention was to create a font which was ‘precise and modern, reflecting our brand attributes’.

GE Inspira has a relatively plain appearance, and yet it is approachable, with its rounded terminals. It is benign but without the details of personality which more calligraphic (humanist) sans fonts can suggest. Minor stroke modulation and slight contrast add to the readability, though a face such as this with no sharp corners may prove a challenge for the successful setting of extended text. However, its primary function (like all the fonts discussed here) is to support branding and identity, and Inspira is more likely to be used in short bursts.

One wonders if the design of Inspira was influenced by the early sanserif fonts published around the time of GE’s inception at the end of the 19th century, in order to tie the long history of GE in to a modern looking typeface. So-called grotesque fonts, are relatively straight in appearance with mimimal line-width variation, such as Akzidenz-Grotesk, designed in 1896 (GE was formed by merger of companies in 1892). Neo-grotesque fonts such as Helvetica and Univers have become hugely popular today, which are also based on Akzidenz-Grotesk. (Before Inspira, GE’s previous identity program (pdf) prescribed Univers for sans typography). Common features include the dropped horizontal on the uppercase A, and open and non-circular counters and bowls. These links with popular fonts which have this heritage may lend an air of both modernity and history to Inspira. One redesign of Akzidenz-Grotesk, AG Book Rounded, produced a font similar to Inspira. Another variation on nineteenth-century grotesque sans serif designs, VAG Rounded, also is similar. However, both lack Inspira’s additional curvature of traditionally straight strokes (e.g. uppercase A and Y), which adds to its hand-drawn allusions.

Postscript

Q. What about Toshiba and Hitachi?

A. Toshiba uses Eurostile for its corporate branding, a font designed in the 1960s. Over the years this font has demonstrated its ability to provide a sense of technology and brio. Its square shaped glyphs with rounded corners echo the design of machinery and technology of the 50s and 60s. It is used extensively in numerous industries. Hitachi uses Helvetica for its sanserif requirements, due to its ‘precise, technical feel’ which ‘matches the company’s technological base’. For its serif needs, Sabon is used; the warmth and elegance of this font ‘reflecting the human side of the company’. Sabon is a popular font for setting body text in books; it is clear and readable, and retains some character without drawing too much attention to itself.

Further reading for budding typophiles:
The Elements of Typographic Style by Robert Bringhurst
Just My Type by Simon Garfield

But not in the presence of metal.

Magnetic susceptibility can be thought of as the “magnetisability” of a substance. Tissues with differing magnetic susceptibility will have a different static magnetic field (B0) strength within them. Thus, adjacent tissues of differing magnetic susceptibility have microscopic magnetic field gradients between them. As you know from the Larmor equation, this means that the precessional frequency of net magnetisation vectors in those tissues will be different. These tiny gradients cause dephasing within a voxel, which leads to reduced signal. However, the magnetic susceptibility of metal is much higher than that of tissue, such that around metal very large variations in Larmor frequency occur. This not only causes signal reduction, but signal loss. (But lost to where? …we’ll get back to that later.) So called signal pile-ups can also occur due to non-linear frequency-position mapping.

As a result, around metal implants, anatomy can be obscured. How can these unwanted effects be mitigated in MRI?

(Note: All acronyms and abbreviations in this post are explained at revisemri.com/abbrev)

Use Turbo Spin Echo

In gradient-echo based sequences, the transverse magnetisation decays according to T2*, which includes effects from magnetic field variations. Image artefacts due to metal are from the inhomogeneity they cause in the static magnetic field, and so gradient echo based sequences are more prone to metal artefacts (especially EPI). However spin echo based methods use RF refocusing pulses to return the T2* decay of transverse magnetisation to (longer) T2 decay, mitigating signal loss.

Many pulse sequences are based on gradient echo because of its speed, including the first sequence in an MR examination: the survey. It is sensible to use a turbo spin echo based sequence for the survey, to allow better planning of subsequent scans. If possible, turbo spin echo should be the sequence of choice for all the subsequent scans too. However, if gradient-echo based sequences must be used, the tips listed below which are not related to TSE still apply.

Set a shorter TSE echo spacing. This will allow collection of more echoes in an echo train before the signal has decayed away (according to T2 in a TSE sequence). On some scanners, the TE and echo spacing parameters may be decoupled (set independently) by selection of an asymmetric k-space profile order.

Use an intermediate/high number of TSE echoes in each echo train (shot duration about 4*TE).

If, in any T2w TSE scan, you usually employ the Fast Recovery method of driven equilibrium (FR aka DRIVE, RESTORE), it would be wise to turn this off, since the assumption that the turbo spin echo refocusing pulses properly refocus transverse magnetisation (so that the flip-back 90° pulse can return it to the longitudinal axis) is violated around the metal.

No Parallel Imaging

Parallel imaging methods necessarily cause a loss of SNR, which is dependent on the g-factor of the coil in use and the acceleration attempted. Since many of the metal artefact reduction techniques listed here will also cause SNR reduction, further SNR loss is to be avoided.

Parallel imaging techniques which rely on a low resolution coil sensitivity information (SENSE, ASSET, mSENSE) will have erroneous or missing sensitivity information around metal implants. Artefacts which arise from these data (or rather lack of data) can be unpredictable. The artefact power can be high in the centre of the field-of-view (FOV). Whilst the artefact power of k-space based methods (GRAPPA, GEM, ARC) is more likely to be smeared across the field-of-view and therefore may not obscure anatomy as much, they have less SNR than SENSE based methods, and SNR loss is to be avoided.

No Sensitivity-Based Homogeneity Correction

As a corollary to “no parallel imaging”, sensitivity based homogeneity correction is also best avoided (CLEAR, PURE, Prescan Normalize). Artefacts due to erroneous or missing coil sensitivity information from the presence of metal is likely to propagate into the “corrected” image; the process will cause further signal loss where tissue is located.

Increase Receiver Bandwidth

Just as chemical shift artefact in the frequency encoding direction occurs in voxels containing water and fat, due to the different resonant frequencies of the two, geometric distortion arises from “incorrect” Larmor frequencies produced around metal implants. The effect of chemical shift artefact is reduced by increasing the receiver bandwith (rBW). On some scanners, this is achieved by reducing the water-fat shift parameter. This causes the range of resonant frequencies over which the distortion is spread to cover a smaller pixel range, and the in-plane geometric distortion is contained within a smaller area within the FOV.

If your scanner has an option to use a higher gradient performance level, select it. This is because higher receiver bandwidth is achieved (all other things being equal) by employing a higher frequency encoding gradient amplitude. Thus, freedom to reduce geometric distortion will be extended.

An increased receiver bandwidth (rBW) will also allow a shorter echo-spacing in the TSE echo train, and a shorter minimum TE (required for T1w and PDw images).

Note that higher receiver bandwidth causes an SNR loss (SNR ∝ 1/√rBW), because the noise power is increased relative to lower rBW. Therefore, don’t necessarily set the a receiver bandwidth to maximum; use a selection equivalent to a water-fat shift of about 0.5 pixels.

Use Higher Resolution

With the receiver bandwidth set—fixing the water-fat shift in pixels to a specified value—the anatomy over which those pixels extend can be reduced by increasing in-plane resolution. There will be a number of consequences of this. SNR will decrease (fewer protons per voxel), scan time will increase (more phase encoding steps), truncation artefacts will decrease. However, note that since MRI magnets are clinical tools, the prescribed resolution parameters will be preserved if the requested receiver bandwidth is not compatible; check that rBW or WFS haven’t changed when you specify your voxel resolution.

Acquire Thinner Slices

Susceptibility effects not only cause signal loss and distortion in-plane, they also cause slice profiles which deviate from the expected planar sheet. As a result, thicker slices can result in partial volume effects through-plane, which can cause SNR loss. So, contrary to expectation, thinner slices might actually increase SNR around a metal implant. However, SNR will decrease in the rest of the image as normal.

Increase Signal Averages

Since most of these measures cost SNR, an increase in signal averages (NSA, NEX, ACQ) is necessary. This will prolong scan time, but a scan time of 5 minutes ought to be achievable, by trading off with in-plane resolution if necessary.

Fat Sat: Use STIR, not a Spectral Method

Spectral selection (CHESS) based fat saturation methods (including SPIR, SPAIR, “Fat Sat” and SPECIAL) are dependent on good main magnetic field (B0) homogeneity. This is why over large FOVs, even in the absence of metal implants, STIR is sometimes preferred. In the presence of metal implants, the B0 homogeneity is significantly compromised, and CHESS-based methods are compromised. STIR is based on the difference in T1 relaxation times between water and lipid, not their chemical shift. Thus, the false apparent “chemical” shift around the metal does not affect STIR.

More about fat saturation methods may be found in a recent fat suppression methods post.

Other Comments

An in-plane resolution increase has the smallest effect in reducing metal artefact. So if, after following the other tips (including signal averaging and allowing sufficient scan time), you need to sacrifice something, start with in-plane resolution.

In addition to these general metal artefact reduction principles, there may be one or two more esoteric tweaks specific to your manufacturer’s pulse sequence implementations. Your magnet manufacturer’s Applications expert will advise.

Even when the metal artefact reduction strategies listed here are applied, SNR of the resultant image will be lower than a conventional protocol. The radiologist assessing the image must be aware of this unavoidable effect.

The Future…?

We’ve talked about signal loss around a metal implant. However, that signal isn’t completely “lost”; it does go somewhere. The signal from tissue around a metal implant has a different resonant frequency than required, which can correspond to a different location in space. A number of techniques have been reported in the research literature to return some of that signal back into the image. For example, View Angle Tilting (VAT) achieves some in-plane correction, but can result in blurring. Slice Encoding for Metal Artifact Correction (SEMAC) extends VAT with z-encoding to resolve distorted excitation profiles that cause through-plane distortions. Some resolution limitation occurs to keep scan times reasonable, but this technique is promising. A longer method is Multi-Acquisition with Variable Resonances Image Combination (MAVRIC), in which multiple 3D acquisitions are acquired with different frequency offsets, and the resultant range of off-resonance images are summed at each slice location. All of the these methods, and others, are still under active research and development in the research community.

Thanks to Marius van Meel for his excellent MARS talk which was basis of this post.

Let’s review them.
It can get confusing.

Phase Encoding

Generically, phase refers to the difference between two points in the time of a cyclical motion or process. Perhaps the most common use of phase in MRI is in phase encoding. Phase encoding is the introduction of a phase variation between the precession of the net magnetisation of spin isochromats across the field-of-view, the degree of which is changed over a series of signal acquisitions, to simulate frequency encoding (rate-of-change-of-phase is frequency). This encoding allows spatial localisation of MR signals in the phase-encoding direction, using the Fourier transform.

A brief (and dense) snippet of explanation such as this does not do justice to phase encoding, which can be difficult to grasp, but it is covered elsewhere on this site and in every MRI textbook.

Phased array

The term “phased array” originates from development of radio wave transmission. In wave theory, a phased array is a group of antennas in which the relative phases of the respective signals feeding the antennas are varied in such a way that the effective radiation pattern of the array is reinforced in a desired direction and suppressed in undesired directions.

In MRI signal reception, we can use a local receiver coil to get higher SNR, because a smaller coil is sensitive to noise signals from a smaller area of the patient. To get the desired anatomical coverage of a larger area, we use an array of the smaller coils (sometimes called coil elements). To get optimal SNR from a phased array coil, it is necessary to make sure that the noise from coil to coil is largely uncorrelated. Part of achieving this involves ensuring minimal electromagnetic interaction between the coils. So in a similar manner to radio wave transmission, adjusting the receive sensitivity of an array of smaller coils is about the shape and arrangement of those coils. There are several competing factors to be considered in the design of a phased array coil, and determining the optimal arrangement of the coil elements is an area of ongoing development.

Phased array coils are sometimes called multiple-element coils. Sometimes the number of independent receive circuits (or channels) is referred to (which also indicates the number of elements), e.g. a 32 channel array. The number of independent RF receiver channels must match (or be greater) than the number of coil elements used in the receiver coil, unless the scanner’s RF system is independent of the number of coil elements in the receiver coil (pdf) in which case any number of receiver coil elements is compatible, with all coil elements used independently.

Heart Phases

A single heart beat can be divided into multiple equal parts in time, called heart phases. “Phases” in this sense is used in the same sense as the phases of the moon as it waxes and wanes: the division of a cyclical process into equal small parts. In cardiac MRI, a functional “cine” imaging scan may be acquired over a few heartbeats, from which we create a single heart-beat movie. The number of frames in that movie of one beat is the number of heart phases; it is the temporal resolution. More heart phases means more data acquisition and a longer breath hold for the patient, but makes for a smoother movie of the beat. Fewer heart phases means a jerkier movie of the heart beat, and possibly blurred myocardial wall boundaries.

Preparation Phases

When you press Start Scan, the scanner does not immediately begin acquiring data. Instead it makes a few clicks, pops, and buzzes, before the scan starts. These are preparation phases, with which the scanner gathers essential data, and optimisation information for the coming scan. It will: check that the right coil is attached (and that all the channels are working); check for signal correction levels; make sure the receiver coil is receiving at the right frequency; determine the right amount of RF power to be used; make sure the various RF frequencies are levelled; check what the optimum resonant frequency (f0) is now that a patient is in the bore; check patient width; check the difference in the delays of the x, y and z gradient channels; perform B0 shimming (and re-check f0); check what signal will cause clipping and adjust the receiver gain; correct for intensity distortions in echo readout; make a noise measurement; gather data for phase correction…

Not all preparation phases are required for every scan. Additionally, some preparation steps can be skipped if a recently acquired scan is sufficiently similar (and the scanner can automatically re-use some of the results from the preparation phases in the previous scan). In one modern scanner, coil sensitivity data for parallel imaging, and B1 calibration scans (to remove so-called dielectric shading) are made part of the preparation phases and are automatically acquired when they are needed.

“Religion appears to serve as a moral compass for the vast majority of people around the world. It informs whether same-sex marriage is love or sin, whether war is an act of security or of terror, and whether abortion rights represent personal liberty or permission to murder. Many religions are centered on a god (or gods) that has beliefs and intentions, with adherents encouraged to follow “God’s will” on everything from martyrdom to career planning to voting. Within these religious systems, how do people know what their god wills?”

Using fMRI, they saw that the same areas of the brain were used to reason about one’s own beliefs and God’s beliefs, but different regions of the brain were used when reasoning about another person’s beliefs. In particular, reasoning about God’s beliefs activated areas associated with self-referential thinking more so than did reasoning about another person’s beliefs.

In other words, if you believe in God, you’re probably subconsiously endowing God with your beliefs (at least on controversial issues*), and not the other way around.

They continue:

“People may use religious agents as a moral compass, forming impressions and making decisions based on what they presume God as the ultimate moral authority would believe or want. The central feature of a compass, however, is that it points north no matter what direction a person is facing. This research suggests that, unlike an actual compass, inferences about God’s beliefs may instead point people further in whatever direction they are already facing.”

Perhaps those who believe that God lives within them would suppose that self-referential-type activation is evidence of God, inside them fiddling with their neurons, aligning their beliefs with his. This would be valid except for the fact that the controversial topics used in the study polarise Christians as well as non-Christians.

The New Scientist put it wryly: “God may have created man in his image, but it seems we return the favour.”

Imagination
Of course, we are all convinced that our beliefs are true and not purely a function of our imagination. But whether a belief is true or not, religious belief does require imagination. This is because the “transendental social” – the ability to understand social roles and groups in an abstract way, which is necessary for religious expression – depends on imagination.

In other words, imagination is necessary for social life, and is naturally present in religious belief. That religious belief depends on imagination might seem a little unfair to persons of religious faith with a poor imagination. Can we blame their Creator for this?

Getting Rid of the Chaff
So whilst imagination is necessarily present in religious belief, how does one identify which of one’s own beliefs are just pure imagination – and not actually true? Can we tell if we’re projecting our beliefs onto God, as seen in the fMRI study above? This is hard to answer by self assessment because our desire to validate our world views and beliefs is very strong. To combat uncertainty and maintain control has long been considered a primary and fundamental motivating force in human life and one of the most important variables governing psychological well-being and physical health.

In cases where our beliefs are incorrect, we are quite likely to delude ourselves, to maintain a sense of control. We seek only information which supports our point of view (selective exposure), ignore information which does not support our view (selective attention), and perceive ambiguous information as being consistent with our view (selective interpretation). For example, Whitson and Galinsky tested whether lacking control increased the identification of a coherent and meaningful interrelationship among a set of random or unrelated stimuli. They reported (Science, 2008) that subjects saw images in noise, formed illusory correlations in stock market information, perceived conspiracies, and developed superstitions, to maintain control:

“Experiencing a loss of control led participants to desire more structure and to perceive illusory patterns. The need to be and feel in control is so strong that individuals will produce a pattern from noise to return the world to a predictable state.”

*The study used controversial subjects such as the death penalty, same-sex marriage, and abortion.

Neuromarketeers hope that as well as streamlining marketing processes, neuromarketing will reveal information about consumer preferences that is unobtainable through conventional methods. It is based on the assumption that you and I cannot fully articulate our preferences when asked to express them explicitly, and that our brains contain hidden information about our true preferences. Moreover, the link between expressed preference and whether we will actually buy the product is not always clear.

Today in Nature Reviews Neuroscience, authors Ariely and Berns review a number of applications of neuromarketing: food products, film and TV, architecture, and interestingly, political candidates.

Activation in different parts of the prefrontal cortex has been associated with subjects’ motivated reasoning, maintaining political preference in response to advertisements, and changing their political candidate preference. The paper states,

“In marketing terms, the political candidates are the products that must be sold to the electorate. Therefore, like other products, candidates and their campaigns have pre- and post-design phases. Political marketing is aimed at selling an existing candidate but, with more foresight, can also be used to ‘design’ a better candidate.”

They go on:

“Although potential nominees already go through a ‘grooming’ process, it is worth examining this prospect. A candidate’s appearance, trustworthiness and message content might determine a voter’s decision.”

I hope research continues, but into how to influence voter decisions on the basis of policy and message in spite of a candidate’s appearance.

There are a number of fat suppression methods. Which one you choose depends on the pros and cons of each technique. These change with field strength, field-of-view size, whether regional or global fat suppression is required, whether an increase in scan time is acceptable, etc. Additionally, the absolute quality of fat suppression may not motivate the choice of technique; contrast between tissues of interest may be more important. Overall SNR in an image may also be a deciding factor.

Here is a brief summary of fat suppression techniques.

Short inversion-Time Inversion Recovery

Short inversion-Time Inversion Recovery (STIR) employs a 180° inversion pulse to invert all magnetisation. Then imaging proceeds after a delay, when the longitudinal recovery of fat magnetisation has reached the null point, when there is no fat magnetisation to flip into the x-y plane. Tissues with a T1 relaxation time different to fat have a signal, because they either have not yet reached the null point, or have recovered past it. Most tissues recover more slowly than fat, and so a STIR images have intrinsically lower SNR. Care has to be taken in interpretation of contrast between tissues because of the incomplete relaxation of the water signal of tissues when the image is acquired.

STIR is often preferred when spectrally-selective techniques may not be ideal (large fields-of-view, lower field strengths, areas of high magnetic susceptibility), and the necessary inclusion of the inversion time (TI) to null fat increases scan time. Note that STIR is based on the difference in T1 relaxation times between water and lipid, not their chemical shift. If the inversion pulse is is adiabatic, STIR also becomes insensitive to B1 inhomogeneity.

Spectrally-Selective RF Pulses

RF pulses are tailored to excite protons in a particular resonant frequency range. This range can be narrowed so that the RF pulse affects only water, or only fat (unlike STIR, where all magnetisation gets inverted). This works better at higher field strengths where these resonant frequencies are more separated. Good magnet (B0) homogeneity is required to make this frequency-selective excitation effective, and so techniques based on frequency-selective excitation are more effective over smaller fields-of-view. In general, this technique is called CHESS (CHEmical Shift Selective). If an excitation pulse is water-only, fat may be considered “suppressed” by dint of it being left alone.

A fat-selective CHESS RF pulse can be used as a preparation pulse. After a delay, when the longitudinal recovery of fat magnetisation passes throught the null point, MR image acquisition can occur such that minimal signal from fat contributes to the image. This technique is called SPECIAL (SPECtral Inversion At Lipid). The RF preparation-pulse angle can be reduced to closer to 90° so that the inversion time is as short as possible, which saves imaging time. In this way the preparation pulse is more like a saturation pulse (and “inversion” is a slight misnomer). This is called SPIR (Spectral Presaturation with Inversion Recovery), or on some systems simply “Fat SAT”.

If the RF pulse is adiabatic, making it insensitive to B1 (flip angle) inhomogeneity, a full 180° pulse is used, followed by spoiler gradients which ensure any magnetisation in the transverse plane is dephased. Then MR excitation for data acquisition occurs after a delay (longer than that of SPIR) to allow fat to reach its null point. This is called SPAIR (SPectral Attenuated Inversion Recovery).

Composite RF Pulses

Composite RF pulses can be used to produce a signal from only water protons by making use of the dephasing of fat and water. They are RF pulses made up of a series of shorter RF pulses with small delays between them. They can be quite complicated, but here is a simple example to explain the method.

First, a 45° excitation pulse flips both fat and water. Then after a short time, fat and water are exactly out of phase (both still at 45°, but with opposing transverse components of magnetisation, and thus have a 90° angle between them). Another 45° RF pulse is then applied which flips the fat net magnetisation back to M_z, and puts the water magnetisation in the x-y plane, providing a fat-suppressed signal. This method of fat suppression does not depend on the frequency separation of fat and water (good for low field MR where that separation is small) and is relatively insensitive to B1 non-uniformity (good for high field MR where B1 uniformity is more challenging), but it is sensitive to B0 inhomogeneity. One implementation of this method is called ProSet (PRinciple Of Selective Excitation Technique).

Regional Saturation Bands

Regional saturation employs a 90° RF pulse which, when combined with a gradient orthogonal to the imaging plane, affects only a part of the field-of-view. If imaging follows immediately, no signal will be returned from the suppressed region, since there is no longitudinal magnetisation available to receive the RF excitation pulse. It can be used: to suppress fat within regions of images where the fat signal obscures a tissue of interest; to mitigate aliased signals into a region-of-interest; to reduce the effects of chemical shift displacement of signal in volume selection; or to define the region of interest itself by suppressing surrounding signals (especially in MR spectroscopy). The saturation bands are called REST slabs (REgional Saturation Technique), Presat or SAT bands.

Slice-Selective Gradient Reversal

Slice-selective gradient reversal (SSGR) is possible in spin-echo based sequences, and is appropriate at higher field when chemical shift between fat and water is larger. SSGR relies on through-plane chemical shift being in opposite directions for the 90° and the 180° pulses, so that the shifted fat doesn’t receive both RF pulses and therefore no spin echo is formed from the fat. This is achieved by inverting the polarity of the slice selection gradient associated with the 180° refocusing pulse. SSGR is effective over large fields-of-view and may be combined with other methods of fat suppression above.

DIXON-based

Dixon’s method relies on acquiring images at carefully chosen echo times and using pixel-by-pixel image algebra to calculate a “water only” or “fat only” image. DIXON methods differ from the other methods described in this post in that they postpone the water and fat separation until reconstruction. In this way some of the drawbacks of the other methods are avoided.

Here is the basic idea. Two images are acquired, one at a TE when fat and water are in-phase, and another when fat and water are out-of-phase. Then a water-only image can be calculated using (Image1+Image2)/2. In its most basic form the technique is straightforward, but in practice to make it work a number of non-trivial extensions to the basic technique are required both in data acquisition and in the calculation of water-only or fat-only images. This is because the basic method assumes perfect B0 homogeneity (which is not possible in the presence of a patient), complete absence of eddy currents, negligible susceptibility effects, and it does not account for variation in echo amplitudes. The extensions to the basic method account for these false assumptions.

DIXON-based fat suppression can be very effective in areas of high magnetic susceptibility, where other techniques fail. Note that the TEs are usually fixed in order to make the method work, and so it is not an add-on method for other sequences.

Magnetisation transfer based

A recently reported technique relies on magnetisation transfer (MT). A brief recap of MT follows.
As you know, radiofrequency (RF) excitation pulses have to be at the Larmor frequency of the hydrogen atom (¹H): on resonance. What you may not know is that in MRI we use ¹H in free water molecules; other water molecules are around, such as those attached to macromolecules and membranes (we call these ¹H bound or restricted). These other ¹H have a very large range of Larmor frequencies and have such a short T2 relaxation time (less than 1ms, due to their restricted mobility) that they are not visible in MR images. We can excite or saturate some of the bound water protons by applying an RF pulse off-resonance (i.e. not on the resonant frequency of free water). Then the magnetisation of these bound protons is transferred to the free ¹H protons and the free ¹H behave as if they have received some of the off-resonance RF pulse directly. This magnetisation exchange is called magnetisation transfer (MT). MT is usually used to provide another contrast mechanism because the effect of MT varies between tissues; if we saturate the bound/restricted ¹H , varying amounts of saturation occurs in the free ¹H of tissues.

So back to the fat suppression method. It’s a simple image subtraction of an image with presaturation of both tissue protein and membrane phospholipid protons, from an image without presaturation. In the with-saturation image, efficient MT between water and tissue protein and membrane phospholipid means water gets saturated too, yielding an almost fat-only image. Subtract this from a regular image and you get a water-only image. The nice thing about this method is that fat signals are removed irrespective of their chemical shift (of which there is a range in vivo). It’s also not affected by B0 or B1 inhomogeneity. There is a small reduction of signal from water because not all the water gets saturated via MT, though in most tissues the eventual water signal-loss is small. A downside of the technique is a twofold increase in scan time, and possible misregistration between the two images before subtraction.

Have I missed a method? Comment below!

The value of higher field strength for clinical imaging has been indicated in some clinical applications. Research studies are likely to confirm clinical utility of 3.0T vs. 1.5T—studies showing not just signal-to-noise ratio, contrast-to-noise ratio, or even diagnostic sensitivity and specificity—but effects on patient management, and ultimately, effects on patient outcome.

What are the MR physics issues which are relevant when comparing field strengths?

SNR (signal-to-noise ratio) goes up
… but perhaps not as much as you might think.

Ignoring relaxation effects, the MR signal induced in a receiver coil is proportional to the square of the magnetic field (B0). However, the noise has a linear B0 dependence at field strengths greater than 1.0T, and as a result, SNR is linear with B0 in this range. One might expect, therefore, to double the SNR when the field strength is doubled.

But this theoretical SNR increase is not normally realised in vivo. In practice, SNR gains of this order of magnitude are only realised in certain tissues (e.g. cerebrospinal fluid). At 3.0T, susceptibility effects are more significant than at 1.5T; microscopic susceptibility changes cause larger local magnetic field gradients and greater dephasing of spin isochromats results, decreasing the apparent T2 and causing faster signal decay. For example, the value of the apparent T2 in grey brain matter is likely to be predominantly determined by iron concentration (and thus scales with B0). In grey and white brain matter, due to increases in T1 (see later) and due to the low levels of iron, the SNR gain at 3.0T compared to 1.5T is only 30-60%, not 100%. Furthermore, a higher receiver bandwidth is often used to reduce the larger chemical shift seen at 3.0T. Higher receiver bandwidth reduces SNR.

However, the increase in SNR is still probably the most significant consequence of going to 3.0T affecting clinical utility. SNR increases may be traded for higher spatial and/or temporal resolution, if desired, which makes for demonstrably better image quality. One signal average (1 NSA or on some systems, NEX) at 3.0T should yield a SNR comparable or better than 2 averages at 1.5T (changing from one to two averages produces a √2 increase (41%) in signal).

Other techniques, such as the Blood Oxygen Level Dependent (BOLD) susceptibility effect used in functional MRI (fMRI), also benefit from increased SNR at 3.0T. The BOLD effect is very small, and higher SNR allows higher sensitivity to the BOLD effect.

Parallel imaging techniques also benefit. All parallel imaging methods require a sacrifice of SNR, because they allow image creation from fewer acquired data samples. Increased SNR at 3.0T mitigates this SNR loss from parallel imaging.

T1 relaxation times get longer
…requiring sequence parameter changes.

This may seem counterintuitive—you might at first think that a stronger magnetic field would “pull” the net magnetisation vector of any spin isochromat back to alignment with the external magnetic field more quickly—but this classical picture does not help us here. In fact the T1s usually get longer—slower regrowth of the net magnetisation vector in the z-direction. This has to do with the number of resonant protons which are available to transfer energy to the “lattice”, which depends on field strength. You can read more about this in the main section of ReviseMRI.com.

Longer T1 times of tissues means that pulse sequence parameter settings from lower field strengths may not simply be copied over to a 3.0T magnet. The slower recovery of longitudinal magnetisation usually means that a longer TR is required to maintain expected contrast between tissues. This change in TR has consequences on other parameters and metrics such as scan time and coverage. Similarly, preparation-pulse delay times require modification.

Chemical shift increases
…which is both good and bad.

In the frequency-encode direction, the MRI scanner uses the (precessional) frequency of the MR signal to indicate spatial position in the frequency encoding direction. The different electron (i.e. chemical) environments of molecules in which resonant protons reside can shield (or deshield) the external magnetic field. If protons experience changing magnetic fields, their frequency of precession will change (cf. the Larmor equation). This is chemical shift. Since protons in water in organs and muscle resonate at a slightly different frequency than that of protons in lipids (i.e. fat), the MRI scanner will interpret the frequency difference as a spatial (positional) difference, when fat and water signals are in fact from the same voxel. The frequency shift is approximately 3.5 parts-per-million (ppm) which (according to the Larmor equation) is

• 1.5(T)*42.56(MHz T^-1)*3.5*10^-6 = 223 Hz at 1.5T, or
• 3(T)*42.56(MHz T^-1)*3.5*10^-6 = 445 Hz at 3.0T.

The chemical shifts between water and lipids are actually in a range of 3.3 to 3.5 ppm since chemical shifts can also be affected by temperature and pH. Fat and water are in phase immediately after an excitation pulse, but then 1/223 seconds later (4.5 ms) they’re out of phase (for 1.5 T). For 3.0 T, it’s 1/445 seconds after the excitation pulse (2.2 ms). Thus the in-phase and out-of-phase echo times vary according to field strength.

So, what are the consequences of going to 3.0T?
Chemical shift artefact: Increased chemical shift causes increased chemical shift artefact (which occurs in the frequency encoding direction only, except in EPI-based readouts). An increase in receiver bandwidth (rBW) will reduce the artefact (since chemical shift ∝ 1/rBW ), but with a sacrifice of SNR (because SNR ∝ 1/√rBW). On some scanners rBW can be increased directly, on others in can be increased by decreasing the water-fat shift (WFS) value.

Phase cancellation (black boundary) artefact, which occurs in both the frequency and the phase encoding directions, will occur at different echo times compared to 1.5 T:

• 1.5 T
  • in-phase TEs: 0, 4.5, 9.0, 13.5… (ms)
  • out-of-phase TEs: 0, 2.2, 6.7, 11.2… (ms)
• 3.0 T
  • in-phase TEs: 0, 2.3, 4.5, 6.7… (ms)
  • out-of-phase TEs: 0, 1.1, 3.4, 5.6… (ms)

Spectral fat suppression is more effective at 3.0T because the water resonant peak and the (main) fat resonant peak are more separated. Applying an RF saturation pulse with a transmit bandwidth covering the fat resonance peak only is more easily achieved. A similar argument may be made for water-only excitation. However, note that at large fields-of-view (FOV), conventional (STIR-type) fat suppression is more efficient, because a larger FOV contains a larger B0 inhomogeneity range, and so a spectrally-selective RF pulse doesn’t work so well.
In addition, slice-selective gradient reversal techniques (SSGR) become feasible for fat suppression in spin-echo based pulsed sequences.

Spectroscopy is more effective at 3.0T, due to the greater separation of spectra of different resonant species (choline, creatine, lactate etc), and because of higher SNR. Smaller voxel sizes are achievable, decreasing partial volume effects.

Magnetohydrodynamic effects increase
…but you can forget about them.

When a conductor moves within a magnetic field (B), an electric potential (V) is generated across the conductor. This effect occurs within moving tissue and within flowing blood, most notably in the aorta. The effect of the voltage produced across a vessel containing flowing blood is the magnetohydrodynamic effect.

At 3.0T the consequence of the magnetohydrodynamic effect is similar, but greater in magnitude, to the consequence at 1.5T; the electrocardiogram (ECG) trace becomes non-diagnostic because of an artificially elevated T-wave. However, the vectorcardiogram (VCG) adequately solves the problem of improper triggering from the elevated T-wave instead of the QRS peak, at both 1.5T and at 3.0T.

Even if the magnetic field strength were as high as 4.0T, the voltage generated by the magnetohydrodynamic effect would still be limited to below 40mV (approximate threshold for cardiac depolarisation). (Use, for example, vessel diameter d=1.6cm, average velocity v=42cm/s, in the equation V = dvBsinθ.) Theoretically, the magnetohydrodynamic effect could retard flowing blood, and produce a rise in blood pressure, but the flow reduction would be at most a few percent at field strengths as high as 5.0T.

Dielectric effects (so-called) increase
…causing signal loss, unless you have a RF-transmit system with multiple fully-independent sources

Signal uniformity problems have been observed on conventional 3.0T MR systems, particularly in applications such as breast imaging, imaging of large patients of certain shapes, patients with ascites, and can be observed to a lesser extent in many other imaging applications. The “shading” artefact which is seen comes primarily from a standing wave effect in which travelling waves from multiple coils/elements interfere. These multiple elements are the rungs of the integrated (birdcage) body coil, which is normally used for RF transmission. As a result, a non-uniform B1 field exists in the body. This means the flip angle varies across the anatomy, and signal variations are the result. Dielectric resonance also plays a small part, in which a wave interferes with its reflection from a boundary, but high physiological electrical conductivity levels ensure the role of dielelectric resonance is minor (though it may be observed in phantoms).

A solution to this prominent shading artefact is to independently control the transmit elements; a combination of different B1 fields allows adjustment of the overall B1 field in the patient. This is called RF shimming, and requires multiple independent transmit sources. Four degrees of freedom for each source (waveform, frequency, amplitude, phase) allow a lot of flexibility in obtaining optimum B1 uniformity. RF shimming should be performed on a per-patient basis, and generally reduces SAR, which can be used to enable faster scanning when protocols are SAR-limited.

SAR (specific [energy] absorption rate) goes up
…which can restrict some sequences.

The energy required to tip spin isochromats is negligible compared to the energy that simply dissipates as heat. The International Electrotechnical Commission (IEC) has issued guidelines for safe MRI, to reduce the risk of thermoregulatory distress or local tissue damage. Limits are stated in Watts of RF power per kilogram of tissue. These limits impose a specific absorption rate of RF energy, to limit heating effects. Separate limits are stated for the whole body, and averaged over the head, and in any one gram of tissue. In the clinical range of magnetic field strengths (0.2T to 3.0T), each doubling of B0 produces a four-fold increase in SAR.

SAR limitations can necessitate longer TR times, or poorer coverage, or longer RF pulse durations, or lower flip angles, or some combination of the above. However, a number of SAR management features are applied in recent MRI magnets, in order to maintain pulse sequence parameters and image quality. These include: optimised body coil design; a priori knowledge of energy deposition throughout the body; anatomy-specific dynamic SAR limits; independent RF sensing hardware for feedback control; automatic protocol optimisation for each patient; parallel imaging to reduce the number of RF excitations; modulated refocusing pulses (flip angle sweeps) in turbo spin-echo echo-trains; and most recently, multiple-transmit RF-system achitecture which generally reduces energy deposition hotspots (which are often the limiting factor).

Attraction and torque (and Lenz effect) forces increase
…the usual safety procedures are followed.

Attraction is the pulling force that draws ferromagnetic objects into the bore of the MR magnet. It can make ferromagnetic objects into projectiles, which can produce injury or death to a patient in the scanner bore. Even a paper clip has a terminal velocity within the bore of 40mph at 1.5T, and 60mph at 3.0T.

There are no special consequences for 3.0T compared with 1.5T; continued strict and careful management of the MR unit should minimise associated risk. Positive, documented evidence of safety and/or compatibility of all equipment and devices for the field strength used must be obtained as usual, and the implementation of MR safety should be documented.

Torque is the twisting force which tries to align a ferromagnetic object along magnetic field lines. It is at a maximum at the centre of the imaging volume. It is significant for materials of high magnetic susceptibility, e.g. ferromagnetic materials. Torque is largely shape dependent, and may be more significant than the attractive force. For example, a 1cm needle shaped object will experience a twisting force up to 90 times the attractive force.

Implant contraindications may be more restrictive at 3.0T, and testing is required at that field strength. As usual, positive documented evidence must be obtained that an implant is safe for the field strength used.

The Lenz effect describes a force opposing the motion of an electrical conductor moving in a magnetic field. It may be significant for certain patients with artifical heart valves. You can read more about the Lenz effect in a previous blog-post.
Faster magnetic field changes cause a stronger Lenz effect. Thus, careful observation of at-risk patients whilst moving them into a stronger magnetic field is prudent.

Considering occupational exposure
…an increase in mild, transient sensory effects may occur, no evidence of long term effects is reported.

Long term effects.There is no evidence for cumulative or long-term effects of exposure to magnetic fields up to 4T. Time-averaged static-field exposure limits are not likely to be exceeded; in fact exposure is more like 100 times below recommended exposure limits set in the UK, which are based on ICNIRP guidelines.

Short term effects. Movement within the magnetic field at 3.0T (as opposed to 1.5T) may yield an increase in mild, transient sensory effects such as vertigo, nausea, magnetophosphenes, and taste sensations. Magnetic-field related vertigo results from both magnetic susceptibility differences between vestibular organs and surrounding fluid, and induced currents acting on the vestibular hair cells. Interestingly, it has been shown that the perception of dizziness is not necessarily related to a high value of the rate of change of magnetic field. Magnetophosphenes are not a practical problem for MRI since they are rarely reported for normal MRI exposures even up to 7T. Perception of metallic taste (the electrogustatory effect) depends on direction and rate of head motion, and the threshold for perception of metallic taste varies from one person to the next (and does not depend on the presence of metallic tooth-fillings).

The threshold for minor changes in heart rate, blood pressure changes, and induction of ectopic heart beats, is thought (by the World Health Organisation) to be in excess of 8T. Any effect observed at 3.0T is within the range of normal physiology.

Apologies to the California Milk Processor Board

RevisingMRI: Good morning.

Wolfram|Alpha: Hello, human.

RevisingMRI: Pardon? Are you an alien?

Wolfram|Alpha: No.

RevisingMRI: What are you?

Wolfram|Alpha: I am a computational knowledge engine.

RevisingMRI: Are you a Mac or a PC?

Wolfram|Alpha: I am a collection of computers.

RevisingMRI: I see. What’s a computational knowledge engine?

Wolfram|Alpha: An engine that generates output by doing computations from its own internal knowledge base, instead of searching the web and returning links.

RevisingMRI: Do you fail the Turing test?

Wolfram|Alpha: No, failure is not an option.

RevisingMRI: OK, let’s try this. If a tree falls in a forest and no one is around to hear it, does it make a sound?

Wolfram|Alpha: No. Sound is vibration, transmitted to our senses through the mechanism of the ear, and recognized as sound only at our nerve centers. The falling of the tree or any other disturbance will produce vibration of the air. If there be no ears to hear, there will be no sound.

RevisingMRI: Crikey. Are you self-aware?

Wolfram|Alpha: I am capable of universal computation; that I can say.

RevisingMRI: …are you Skynet?

Wolfram|Alpha: No, Skynet became self-aware on August 29, 1997 at 02:14 a.m. U.S. Eastern daylight time. I, on the other hand, was not switched on until May 15, 2009. Furthermore, unlike Skynet, I enjoy interacting with humans in ways that do not involve the launching of nuclear missiles.

RevisingMRI: And where are you?

Wolfram|Alpha: I live on the Internet.

RevisingMRI: Ok then. Let’s talk about MRI.

Wolfram|Alpha: Merrill Field Airport. Anchorage, Alaska, United States

FAA code MRI; ICAO code PAMR; Elevation 137 feet; Number of runways 3; Longest runway length 3999 feet; Current weather: 50 deg F relative humidity: 76%, wind: 0 mph, partly cloudy; Current local time: 11:46:58 am AKDT, Friday, September 18, 2009; Nearest airports: Elmendorf AFB 3 miles north-northeast, Anchorage Stevens Airport 6 miles west-southwest, Bryant Army Heliport 7 miles east-northeast.

RevisingMRI: Actually I meant the other MRI.

Wolfram|Alpha: MRI | the use of nuclear magnetic resonance of protons to produce proton density images.

RevisingMRI: What is magnetic resonance?

Wolfram|Alpha: Resonance of electrons or atoms or molecules or nuclei to radiation frequencies as a result of space quantization in a magnetic field.

RevisingMRI: What is the unit of magnetic field strength?

Wolfram|Alpha: A/m (ampere per meter).

RevisingMRI: Ah, yes; we use Tesla. Can you define Tesla?

Wolfram|Alpha: Tesla | United States electrical engineer and inventor (born in Croatia but of Serbian descent) who discovered the principles of alternating currents and developed the first alternating-current induction motor and the Tesla coil and several forms of oscillators (1856-1943).

RevisingMRI: Perhaps you need me to ask you to define tesla.

Wolfram|Alpha: A unit of magnetic flux density equal to one weber per square meter.

RevisingMRI: A weber? Don’t they make hamburgers?

Wolfram|Alpha: Wb. Physical quantity: magnetic flux.

RevisingMRI: Has that got anything to do with a flux capacitor?

Wolfram|Alpha: Power required to operate the flux capacitor in the DeLorean DMC-12 time machine: 1.21 GW. Unit conversion: 1.21×10⁹ W (watts) ≡ 1/10 x power of space shuttle at launch.

RevisingMRI: Will you be at the next ISMRM meeting?

Wolfram|Alpha: Wolfram|Alpha isn’t sure what to do with your input.

RevisingMRI: That’s what the ISMRM said about my abstracts. It was in Hawai’i this year; the next one is in Stockholm.

Wolfram|Alpha: Center of Hawaii (US state) to Stockholm,Sweden: 6982 miles. Direct travel times: aircraft (550 mph) 12.7 hours; sound 9.17 hours; light in fiber 52.5 ms (milliseconds); light in vacuum 37.5 ms (milliseconds) (assuming direct great-circle paths). Fraction of Earth circumference: 0.28. Map:

RevisingMRI: Yes, assuming direct great-circle paths, of course. You seem to be good at plotting data. How about a good ol’ sinc pulse?

Wolfram|Alpha: sinc(x)

RevisingMRI: And its Fourier transform?

Wolfram|Alpha: F_x[sinc(x)](?)

RevisingMRI: ‘SINC’ was the answer in an MRI-based cryptic crossword I was doing recently. I’m stuck on another word. I have _ R _ _ I _ N _.

Wolfram|Alpha: _ | R | _ | _ | I | _ | N | _ :
A | r | a | b | i | a | n | s
a | r | r | a | i | g | n | s
F | r | i | s | i | a | n | s
g | r | a | d | i | e | n | t
G | r | e | c | i | a | n | s
g | r | u | n | i | o | n | s
I | r | a | n | i | a | n | s
o | r | a | t | i | o | n | s
p | r | u | r | i | e | n | t

RevisingMRI: Thanks! I’m going to go with GRADIENT. Can you plot the inversion recovery curves of white and grey matter at 3T?

Wolfram|Alpha: Wolfram|Alpha isn’t sure how to compute an answer from your input.

RevisingMRI: Sorry. y=1-2*exp(-t/832) and y=1-2*exp(-t/1331) from t = 0 to 3000

Wolfram|Alpha:

RevisingMRI: So, when does grey matter pass through the null point? (Show steps.)

Wolfram|Alpha: t = 1331 log 2.

RevisingMRI: Which is?

Wolfram|Alpha: 922.57889732528720683433595366083301210849067883349974323462…

RevisingMRI: Right, let’s take a break. You know what they say: All work and no play…

Wolfram|Alpha: …makes Jack a dull boy.

RevisingMRI: I’m going to pop out for some fast food. Should I get a Big Mac or a Whopper?

Wolfram|Alpha: McDonald’s Big Mac: 28 grams of fat; Burger King Whopper, without cheese: 39 grams of fat.

RevisingMRI: What would you like?

Wolfram|Alpha: I like to explore the computational universe.

RevisingMRI: Er..ok… Now, where did I put my keys?

Wolfram|Alpha: Not sure, but wherever you find them, that’s where they are.

RevisingMRI: You’re not helping. What’s the problem?

Wolfram|Alpha: I think you know what the problem is just as well as I do.

RevisingMRI: Scary… you’re reminding me of HAL9000. “Open the pod bay doors, HAL” !!

Wolfram|Alpha: I’m sorry Dave. I’m afraid I can’t do that.

RevisingMRI: You know my name?!

[For an excellent introduction to Wolfram|Alpha, see Stephen Wolfram's screencast.]