1 . LA EVICCIÓN Y LOS VICIOS REDHIBITORIOS
A. EL SUPUESTO DE HECHO
2. EL CONOCIMIENTO
The permanent magnet is the most familiar type of magnet. This is the type of magnet used to demonstrate magnetic fields in science classes and to fix paper to refrigerators.
Permanent magnets are components of com-passes, motors, and audio speakers. They are inexpensive and widely used for simple applications.
Permanent magnets occur naturally, or they can be synthesized.
The earliest commercial magnets were made of iron and called ferrite magnets. In the 1930s, an alloy called alnico (aluminum, nickel, and cobalt) was developed with a slightly higher magnetic field than ferrite magnets. Alnico magnets have been made with ever-increasing magnetic field intensities. More recently, rare
earth magnets have been introduced, which have even higher magnetic field intensity (Figure 11-2).
Although magnetic field strengths of up to approximately 1.2 T can theoretically be achieved with permanent magnets, field strengths of only about 0.3 T are practical for whole-body MRI systems. Figure 11-3 shows a typical permanent magnet MRI system.
This design is often called an open MRI sys-tem because it enables parents to remain with their child during imaging. Open permanent magnet systems also make claustrophobic or anxious patients more comfortable. The mag-netic field is typically produced by individual brick-size ferromagnetic ceramic materials that are rendered magnetic by charging them in the field of an electromagnet (Figure 11-4).
Once magnetized, these bricks are then care-fully oriented into an array, up to 1 m on a side, containing two to five layers. The fabrication of such a large magnet made from smaller mag-nets is not a trivial task. The forces exerted are enormous, and if one brick is positioned incor-rectly, contrary magnetic fields can result and cause the whole assembly to fragment violently.
A variety of magnetic bricks are shown in Figure 11-5. Two assemblies are positioned opposite one another at a distance appropriate for head (50 cm) or whole-body (100 cm) imaging. A typical permanent magnet design is shown in Figure 11-6. In such a design the four corner posts are iron and provide a return path for the magnetic field.
A pole face is positioned on each magnet assembly (Figure 11-7). These pole faces are Box 11-1 Types of Magnets Used
to Generate B0
MRI, Magnetic resonance imaging.
MRI B0MAGNETS
Permanent Electromagnets
Resistive Superconducting
TABLE11-1 Characteristic Features of Magnetic Resonance Imaging Systems
Resistive Superconducting Permanent Magnet Electromagnet Electromagnet
B0field intensity 0.1-0.3 T 0.2 T 0.5-4 T
Power requirements Low Very high Low
Cooling requirements None Water Cryogenic
Magnetic fringe field None Modest Strong
carefully machined iron slabs designed to help orient and shape the B0field and to increase its homogeneity within the imaging volume.
Often there are adjusting screws or other mechanical shimming devices to further refine the homogeneity of the magnetic field after the imaging system is installed in a prepared site.
Permanent magnets are shimmed with a pole face.
An iron yoke is in physical contact with each permanent magnet, and this contact serves
three purposes. First, it provides a mechanical frame for assembly and stability. Second, it confines the fringe magnetic field by
concen-Rare earth
Relative magnetic field intensity
Alnico Ferrite
1910 1920 1930 1940 1950 1960 1970 1980 1990 2000
Figure 11-2 Developments in permanent magnet design have caused magnetic field inten-sity to increase over the years.
Figure 11-3 A permanent magnet magnetic reso-nance imaging system featuring an open design pro-vides easy access and comfort for claustrophobic patients. (Courtesy Hitachi Medical Corp.)
Figure 11-4 A method for rendering ceramic bricks magnetic with an electromagnet.
trating the lines of the magnetic field in this iron yoke. Finally, by containing the fringe magnetic field, the yoke also intensifies B0 in the imaging aperture. Without an iron yoke, B0 strength of such an assembly would be much less (Figure 11-8). The yoke is usually made of soft iron laminated and bolted together like a transformer core.
Table 11-2 presents the principal character-istics of a permanent magnet MRI system. The principal advantage of a permanent magnet MRI system is the insignificant fringe magnetic field, which for any MRI system must be no greater than 0.5 mT in any controlled area.
This level is chosen out of consideration for patients with cardiac pacemakers; it is not haz-ardous to others.
The 0.5-mT field is within a few centimeters of the permanent magnet gantry because of the mass and design of the iron yoke. Other advan-tages of a permanent magnet MRI system include low electric power consumption and the absence of a cooling system.
The principal disadvantage of a permanent magnet MRI system is the limited B0intensity.
This places some restrictions on the type and
complexity of imaging allowed. Other disad-vantages include the relatively poor magnetic field homogeneity, usually about 20 parts per Figure 11-5 Individual magnets of varying sizes and shapes are assembled to produce the strong magnetic field of a permanent magnet magnetic resonance imaging system. (Courtesy Sumitomo Special Metals.)
Figure 11-6 Typical permanent magnet design for imaging. (Courtesy Sumitomo Special Metals.)
million (ppm), and the excessive weight. Poor magnetic field homogeneity results in reduced spatial and contrast resolution. The weight of a permanent magnet MRI system limits its use to fixed sites.
Permanent magnets are attractive for low magnetic field imaging applications because of their minimal fringe magnetic fields, low power requirement, and open architecture.
However, the use of permanent magnets in Pole face
Adjusting screw
Figure 11-7 Permanent magnets are shimmed by positioning precisely machined pole faces.
B0
B0
Figure 11-8 The presence of an iron yoke (A) with a permanent magnet imaging system intensifies the B0field in comparison to one without (B). The yoke provides a return path for the lines of the magnetic field.
A B
clinical MRI is limited. This is because the maximum B0intensity is low. Permanent MRI systems compromise the ability for some clini-cal applications, such as echo planar imaging.
Although the fringe magnetic fields tend to be small, the weight of a well-designed permanent magnet MRI system can approach 90,000 kg (approximately 100 tons). This imposes signifi-cant mechanical considerations on the chosen location and usually excludes all but ground floor siting.
There is a fundamental difference between a permanent magnet MRI system and one of
electromagnet design. The B0field of a perma-nent magnet imaging system is vertical (Figure 11-9). Therefore the Z-axis is vertical rather than horizontal as in most superconducting electromagnet imaging systems. The long axis of the patient in a permanent magnet imaging system is the X-axis, and the lateral direction is the Y-axis. This corresponds to the axis identi-fication in vector diagrams.
ELECTROMAGNETS
Electromagnets make up the second type of primary magnet system for imaging. It is use-ful to discuss resistive electromagnets and superconducting electromagnets separately because they have significantly different oper-ating characteristics. However, the physics of image production is exactly the same regard-less of magnet type, especially when compar-ing equal B0. Neither the patient nor the secondary magnets can identify the origin of the B0. There are no characteristic image dis-tinctions.
Resistive Electromagnets
Resistive electromagnets are making some-thing of a comeback, having nearly disap-peared from the commercial MRI scene. The B0 in a resistive electromagnet imaging system is produced by a large, classical electromagnet.
All of the early MRI investigations were con-ducted with such resistive air-core electromag-nets. The two outside coils of the classical four-coil design are smaller in diameter, ren-dering the B0more uniform in the imaging vol-ume between the two large inside coils (Figure 11-10).
One variation of a resistive electromagnet MRI incorporates the design shown in Figure 11-11. This arrangement results in a vertical B0 field, which can be intensified by coupling to a permanent magnet.
The most common design of solenoid resis-tive electromagnets uses aluminum strips wound spirally around a tube in several thou-sand layers. The advantage in cost and weight TABLE11-2 Characteristics of a
Permanent Magnet Magnetic Resonance Imaging System
Feature Value
Magnetic field (B0) Up to 0.3 T Magnetic field homogeneity 10-50 ppm
Weight 90,000 kg
Cooling None
Power consumption 20 kW
Distance to 0.5-mT fringe field <1 m
Z
X Y
Figure 11-9 The long axis of the patient is the Y-axis in a permanent magnet magnetic resonance imaging system corresponding to the axis identifica-tion of vector diagrams.
dictates aluminum as the preferred resistive electromagnet material. The mass density of aluminum is approximately one third that of copper; however, it has only about 60% the conductivity of copper. Table 11-3 gives the principal characteristics of a resistive electro-magnet MRI system.
The maximum field strength of the resis-tive electromagnet is approximately 0.3 T.
Resistive electromagnet imaging systems have some advantages. They are less expensive than superconducting electromagnet imaging sys-tems because they operate at lower field strengths and they do not require the precision and homogeneity of a superconducting imag-ing system. Magnetic field homogeneity of 10 to 50 ppm exists for most resistive MRI
sys-tems. Shimming this type of magnet is some-what less difficult than shimming a supercon-ducting system.
Because such a resistive electromagnet is readily brought up to designed magnetic field strength, it is just as easily turned off. This removes the hazard of ferromagnetic projec-tiles during nonimaging time. It is a simple matter to remove metallic objects that become attracted to and stuck in the magnet.
Siting the resistive electromagnet imaging system is easier than siting a superconducting electromagnet in terms of fringe magnetic fields. A resistive electromagnetic imaging sys-Figure 11-10 A resistive electromagnet usually has four separate coils to intensify B0and make it uniform.
B0
Figure 11-11 This resistive electromagnet config-uration produces a vertical B0field.
TABLE11-3 Characteristics of a Resistive Electromagnet Magnetic Resonance Imaging System
Feature Value
Magnetic field (B0) Up to 0.3 T Magnetic field homogeneity 10-50 ppm
Weight 4000 kg
Cooling Water, heat
exchanger
Power consumption 80 kW
Distance to 0.5-mT fringe field 2 m
tem also weighs less than permanent magnet imaging systems (4000 kg versus 90,000 kg).
For these reasons, siting the resistive electro-magnet imaging system is often the simplest.
The principal disadvantage of this type of imaging system is electric power consumption.
A resistive electromagnet imaging system is the most power hungry of the three. A 0.2-T imaging system may require 60 to 80 kW, and this is a continuous power drain when the magnet is on. In addition, requirements for cooling the magnet must be met. Resistive electromagnets are water cooled, with a closed primary loop communicating with a secondary single pass system through a heat exchanger.
Superconducting Electromagnets
Electromagnets are created by conducting elec-tric current through coiled wire. MRI systems based on superconducting electromagnet tech-nology have the principal characteristic of high B0magnetic field strength.
Most superconducting clinical imaging sys-tems operate at 0.5 T, 1.0 T, or 1.5 T.
Superconducting imaging systems at 3 T and 4 T are in clinical operation at many sites, but these are specialty systems. Superconducting electromagnets used for analytical spec-troscopy and high-energy physics now exceed 14 T. These systems can achieve such high fields because of their high electric current and small bore size.
High field superconducting magnets have relative advantages and disadvantages when compared with low field systems. Higher B0 requires more intense gradient magnetic fields, resulting in broader RF bandwidth.
Longitudinal relaxation time (T1) increases with increasing B0and requires relatively longer repetition time (TR) for the same T1 weighting (T1W). B0has no effect on transverse relaxation time, T2, at field strengths below approximately 2 T. At higher B0, T2 decreases slightly.
The precessional frequencies of fat and water protons are separated by approximately 3.5 ppm. At 0.3 T this amounts to only 45 Hz and is not noticeable. At 3 T the chemical shift
is 450 Hz and results in a distinct artifact.
Artifacts caused by magnetic susceptibility, blood flow, and patient motion are more severe at high B0.
The most common magnets today use super-conducting technology to achieve intense, highly homogenous magnetic fields. Table 11-4 presents the principal characteristics of a clinical superconducting electromagnet MRI system.
Superconductivity
At room temperature, all materials resist the flow of electric current. Superconductivity is the property of some materials that allow them to conduct electricity with no resistance. Such materials reach superconductivity as tempera-ture is lowered to a critical temperatempera-ture (Tc).
This critical temperature varies with each superconducting material.
Superconductivity means that once the electric current begins to flow, it will flow indefinitely.
All clinical superconducting electromagnets use niobium-titanium (NbTi) alloys that have a critical temperature of approximately 9 Kelvin (K). The way to achieve such a low tempera-ture is to immerse the conductor in liquid helium.
TABLE 11-4 Characteristics of a Clinical Magnetic field homogeneity 0.1-5 ppm
Weight 10,000 kg
Cooling Cryogenic
Power consumption 20 kW
Distance to 0.5-mT fringe field 10 m
Liquid helium vaporizes at 4 K and there-fore easily maintains the NbTi conductor below its critical temperature. Liquid nitrogen, which vaporizes at 77 K, was used to insulate the liquid helium. Liquid nitrogen is seldom used in current cryostats. Liquified gases that are used to keep the conductors cold, such as liquid helium and liquid nitrogen, are called cryogens.
The container housing the superconduct-ing wire and the cryogens is a cryostat.
Superconductivity is one of the four electrical states of matter (Box 11-2). In this state, elec-trons flow in a conductor and experience no resistance. As the temperature of a conven-tional electrical conductor such as copper is lowered, its resistance decreases (Figure 11-12).
Theoretically the decrease is linear to the ori-gin, so its value is 0 at 0 K.
A superconductor such as NbTi has higher resistance at room temperature than copper and is therefore a poor conductor at room tem-perature. Its resistance decreases linearly with temperature just as copper. However, when it reaches its Tc, its resistance immediately drops to zero.
There is much scientific investigation of superconductivity at this time that may pro-foundly affect the future of MRI. In 1987 Bednorz and Müller won the Nobel Prize in physics for their work on high temperature superconductivity in a new class of materials.
They showed that the state of superconductiv-ity could be made to exist in some exotic mate-rials at temperatures above 20 K. The matemate-rials showing the most promise are compounds of lanthanum, barium, copper, and oxygen.
Since then, other investigators have demon-strated the state of superconductivity at even higher temperatures (Figure 11-13). If this research results in superconducting wire above the boiling point of nitrogen, 77 K, the cost of MRI systems will be greatly reduced. The design and production of the cryostat would be greatly simplified to accommodate the cheap and readily available liquid nitrogen.
There are several advantages to supercon-ducting MRI magnets. The high magnetic field intensity is essential if spectroscopy is planned; however, the tolerances on field homogeneity are more restrictive at such high B0intensities.
Higher B0intensity is desirable because the increased magnetic resonance (MR) signal from such imaging systems results in higher signal-to-noise ratio (SNR), allowing shorter examination times and fewer motion artifacts.
In addition, because of the increased SNR, higher spatial and contrast resolution images are obtained with these systems.
The field of the superconducting electro-magnet can also be homogenized or shimmed to a degree not achievable with the other mag-net systems. The solenoidal design of
super-Liquid He
Liquid He TEMPERATURE (K) 200
Figure 11-12 The electrical resistance of a con-ductor (Cu) and a superconcon-ductor (NbTi) as a func-tion of temperature. Tc, Critical temperature.
Box 11-2 The Four Electrical States of Matter
1. Insulator 2. Semiconductor 3. Conductor 4. Superconductor
conducting magnets makes them inherently homogeneous. Shimming is important for small field of view (FOV) imaging, optimal fat suppression, fast imaging, and spectroscopy.
Homogeneity of less than 1 ppm over a 40 cm FOV is obtainable with superconductive shim-ming. At 1 ppm, a 1-T imaging system has a B0 of 1 T ± 1 µT (see Chapter 12).
The principal disadvantage of these systems is their intense fringe magnetic field, which can compromise site selection. At a magnetic field strength of 1 T, the 0.5 mT (5 gauss) fringe magnetic field associated with pacemaker exclusion extends some 10 m in all directions.
The extent of the fringe magnetic field may be reduced by passive or active shielding.
Passive shielding consists of placing ferromag-netic material in the walls of the examination room or around the magnet (Figure 11-14).
With a 2-cm ferromagnetic shield, the 0.5-mT isomagnetic line of a 1-T imaging system can be reduced to a distance of perhaps 5 m from 10 m. However, this passive shielding is heavy (as much as 9000 kg) and expensive but quite standard for current magnets.
Active shielding uses a second set of super-conducting coils positioned outside the pri-mary coils but still inside the cryostat with electric current of opposite direction of the pri-mary current. This produces a magnetic field that counteracts the primary magnetic field, resulting in reduced fringe magnetic field out-side the magnet. Although active shielding is expensive and may not appreciably change the weight of the system, it is generally standard now.
The effect of other magnetic fields and fer-romagnetic objects on the fringe magnetic field of actively shielded magnets is poorly under-stood and unpredictable. Figure 11-15 shows the approximate distance to the 0.5-mT exclu-sion line for a 0.5 T, 1 T, 1.5 T, and 2 T unshielded magnet and for a magnet passively or actively shielded.
A new clan of superconducting electromag-nets has been developed to compete with the
“open architecture” of permanent magnet imaging systems. Such imaging systems are designed around two magnets stacked so that the B0 field may be 1.0 T with an imaging
1900 1920 1940 1960 1980 2000
Year
Temperature (K)
?
Figure 11-13 Recent years have shown a dramatic rise in the initial temperature for super-conducting materials.
Figure 11-14 A passively shielded superconducting imaging system. (Courtesy Siemens Medical Solutions, Malvern, PA.)
1m2
Unshielded Passive Shield
Active Shield
1,0T
1,5T 0,5T
2,0T
-9 -8 -7 -6 -5 -4 -3 -2 -1 0 1 2 3 4 5
-7 -6 -5 -4 -3 -2 -1 0 1 2 3 4 5 6
6 7 8 9
Figure 11-15 Distance to the 0.5-mT fringe magnetic field line for several imaging field strengths, both unshielded and shielded.
aperture of 50 cm. If actively shielded, the 0.5-mT fringe magnetic field is restricted to a 4-m distance.
Another disadvantage of the superconduct-ing magnet is the need for cryogenic gases.
Maintenance of the NbTi superconducting wire in a superconducting state requires an envi-ronment below its critical temperature of 9 K.
The NbTi wire is wound on aluminum formers that are immersed in liquid helium, which vaporizes at 4 K.
External forces, principally in the form of thermal radiation, will cause the liquid helium to heat and vaporize over time. Because liquid helium is expensive, its vaporization is reduced by surrounding the helium compart-ment with concentric insulating compartcompart-ments (Table 11–5).
In some magnets, the compartment contain-ing the coils and liquid helium is separated from an outer vessel containing liquid nitrogen by a vacuum shield. Liquid nitrogen, which vaporizes at 77 K, is used instead of liquid hydrogen (20 K) or liquid neon (27 K) because it is less expensive. Furthermore, because of its high boiling point, liquid nitrogen also vapor-izes more slowly.
Temperature Scales TC = 5/9 (TF- 32) TF = 9/5 TC+ 32 TK= TC+ 273
where the subscripts C, F, and K refer to Celsius, Fahrenheit, and Kelvin, respec-tively.
The liquid nitrogen is also protected by a vacuum shield from the environment. This arrangement of multiple thermal compartments is shown in Figure 11-16. The whole assembly is approximately 3.0 m in diameter, but because of the multiple thermal layers, the bore of the assembly is only approximately 1 m.
Advanced magnet design eliminates the nitro-gen compartment in favor of superior vacuum
compartments with liquid helium–replenishing devices called cryogenerators. Figure 11-17 shows an actively shielded 1.5-T imaging sys-tem with a cryogenerator.
These several thermal shields are designed to maintain the superconducting condition of the electromagnet. Despite rather rigorous designs, both liquids tend to vaporize, the nitrogen more readily than the helium. Vaporization rates of approximately 1.0 l/hr for nitrogen and 1.0 l/day for helium are experienced in some
These several thermal shields are designed to maintain the superconducting condition of the electromagnet. Despite rather rigorous designs, both liquids tend to vaporize, the nitrogen more readily than the helium. Vaporization rates of approximately 1.0 l/hr for nitrogen and 1.0 l/day for helium are experienced in some