Medical Imaging Technology

Perfusion MRI

2015-05-05T07:57:00.001-07:00

Perfusion MRI

Perfusion

Perfusion is a fundamental biological function that refers to the delivery of oxygen and nutrients to tissue by means of blood flow.

Perfusion normally refers to the delivery of blood at the level of capillaries and measures in ml/100gm/min.

Perfusion is closely related to delivery of oxygen and other nutrients to the tissue.

Two main perfusion MRI approaches have been developed:

A. Those using exogenous contrast agent

B. Those without using contrast agent

The first group of technique include Dynamic Susceptibility Contrast (DSC-MRI) and Dynamic Contrast Enhanced (DCE-MRI).

The second group relates to Arterial Spin Labeling (ASL).

DSC-MRI Principle

DSC-MRI is one of the exogenous contrast based method and relies on IV injection of paramagnetic contrast agent, Gd-DTPA.

This technique utilizes very rapid imaging to capture the first pass of contrast agent, k/a Bolus Tracking MRI.

After the bolus of contrast agent is injected, hemodynamic signals of DSC-MRI depends on T2 or T2* relaxation time and transiently decrease because of increasing susceptibility effect.

DSC-MRI is based on susceptibility changes after injecting contrast agent.

The contrast agent is a paramagnetic material, which distorts the magnetic field and reduces T2 around the vessel because of an increased susceptibility effect.

Two compartments must be considered: intravascular (IV) and extravascular(EV).

When the tracer remains IV, the compartmentalization of contrast agent creates strong, microscopic susceptibility gradients which extend beyond the vessel size.

The contrast agent modifies the blood T2/T2* relaxation rates (R2/R2*).

The changes of T2/T2* (∆R2*) is the subtraction of R2* b/w post and pre-contrast injection.

DSC-MRI Parameter

Here, a bolus of Gd-based contrast agent administered as a short venous injection of 10-15 sec, by the time it reaches the ROI (brain), creating a signal dip of about 10-20 sec or longer.

To faithfully record the tracer conc during this passage, images must be acquired at a rate much faster than the time it takes the bolus to pass through the tissue (i.e. MTT), which is usually of order of several sec.

Adequate coverage (whole brain) with T2*WI at a time resolution TR<2sec needs rapid imaging seq like EPI.

TE is chosen long enough to produce sufficient CNR due to susceptibility effects, but not long enough to dephase all signal during the max contrast agent conc.

A relatively large FA is used, although not long enough to introduce unwanted T1 contamination.

DSC-MRI Signal Dynamics

Serial images are acquired before, during and after injecting the contrast agent.

While passing through the microvasculature, a bolus of contrast agent produces decrease in SI.

The time course images can be divided into 3 stages: the baseline, the first passage of bolus and the recirculation period.

DCE-MRI

This is the other exogenous contrast based method.

After the bolus of contrast agent is injected, hemodynamic signal of DCE-MRI depend on T1 relaxation time, and increase because of T1 shortening effect associated with paramagnetic contrast agent.

DCE-MRI uses rapid and repeated T1W images to measure the change in signal induced by paramagnetic tracer in the tissue with change in time.

In this method, contrast agent is also IV injected to generate bolus.

Shortening of T1 relaxation rate caused by contrast agent is the mechanism of tissue enhancement (so called T1 or relaxivity based method).

This is usually scanned with T1W in 2D or 3D.

GE methods (FLASH) T1W high resolution isotropic volume acquisition are used.

For quantitative DCE, a pre-contrast T1 value for each voxel is often performed.

These values are used to calculate the post injection T1 values for each voxel.

DCE-MRI Parameter

DCE-MRI is scanned with very short TR and TE to generate T1W images.

FA is also small due to short TR.

With current MRI technique, the volume scan time is usually b/w 5-10sec.

For breast and large FOV, it may extend upto 20sec.

So, it may require compromise in terms of spatial resolution, temporal resolution and coverage.

DCE-MRI Signal Dynamics

The time of enhancement is related to the changes, which depend on the physiological parameters of microvasculature in the lesion and on the volume fractions of the various tissue compartments.

For a bolus inj, there is always an initial rise in its conc in the plasma and possibly some leakage into the interstitium for the duration of the inj.

Afterwards, the plasma conc continuously decreases because of diffusion into the body and clearance through the kidneys to the urine.

The dynamic acquisition pattern is similar to that of DSC-MRI.

Images are acquired before, during and after injection of contrast agent.

Most benign and malignant lesions show signal enhancement in first few min after bolus administration.

The normal tissue may also show enhancement.

ASL

ASL gives absolute values of perfusion of tissue by blood.

This technique utilizes arterial water as an endogenous diffusible tracer which is usually achieved by magnetically labeling incoming blood.

ASL is completely non-invasive, using no injection of contrast agent or ionizing radiation and is repeatable for studying normal or abnormal physiology and its variation with time.

ASL required the subtraction of two images, one in which the incoming blood has been magnetically labeled (k/a label or tag image) and the other in which no labeling has occurred (k/a control or reference image).

The signal diff is the ASL signal and which removes the static tissue signal.

ASL signal is approx <1% which makes SNR.

Types of labeling methods

Continuous ASL (PASL)

Pulsed ASL (PASL)

Velocity Selective ASL (VSASL)

Vascular Territory Imaging (VTI)

CASL

In CASL, the magnetization of the arterial blood flowing through a major artery is continously labeled (usually by inversion) using RF pulses.

It is based on steady state approach.

A continous RF field is applied for a few sec along with a field gradient.

PASL

PASL involves a relatively short RF pulse, which results in labeling (usually inversion) of the blood in a large region adj to the imaging volume.

1. Signal alternating with alternating radio frequency (STAR)

2. Flow alternating inversion recovery (FAIR)

3. Proximal inversion with control for off-resonance effects (PICORE)

4. Transfer insensitive labeling technique (TILT)

VSASL

VSASL does not label the blood based on its spatial location (as in CASL and PASL) but rather does so based on its velocity, e.g. using binomial pulses to selectively label blood flowing below a certain velocity.

This labeling is not spatially selective, but can be made sensitive to different blood velocities.

ASL Parameter

ASL contrast is related to the preparation module (blood labeling), but does not rely on T2/T2*/T1 contrast of the acquisition module.

Short TE to maximize SNR

Long TR to allow the labeling blood to reach an imaging plane.

Large FA is used.

Long scan duration due to NSA around 40.

ASL Signal Dynamics

In routine PASL and CASL, the signal at a single TI or post labeling delay of 1.5-1.8 sec is often used.

Two labeling times (Quantitative Imaging of Perfusion using a Single Subtraction- QUIPSS I/II) are used to minimize the definition error of bolus width.

The ASL signal curve also shows 3 phases.

Summary

DSC-MRI is good for quick measurement of transit time, whole brain coverage and fast scan time.

DCE-MRI is good for measurement of BV, perfusion permeability and capillary permeability and for reducing image artifacts.

ASL is good for BF measurement, repeatable studies and provide unique opportunity to provide CBF information without inj of contrast agent.

However, bolus methods with inj of contrast agent provide better sensitivity with higher spatial resolution.

References

Perfusion Magnetic Resonance Imaging: A Comprehensive Update on Principles and Techniques, Korean Journal of Radiology, July 2014.

Perfusion MRI: Technical Aspects, 2009 Lecture.

Perfusion Imaging by Monil Shah.

Dynamic Contrast Enhanced MRI by Sungheon Gene Kim, PhD, NYUSOM/Radiology/CBI.

MR Perfusion Imaging Techniques and Applications, Chen Lin, PhD, Indiana University School of Medicine & Clarian Health Partners.

ASL Basics I, University of Navarra, Pamplona, Spain.

Alternatives to BOLD for fMRI, Harvard-MIT Division of Health Sciences and Technology.

Perfusion MRI: The Five Most Frequently Asked Technical Questions, AJR Am J Roentgenol, 2013.

Contrast-Enhanced MR Perfusion Imaging, University of Wisconsin, Madison, RSNA Chicago, 2008.

Advanced MRI Techniques (and Applications), University of California, Los Angeles, January 2012.

Noncontrast Perfusion MR Imaging with Arterial Spin Labeling (ASL) of cerebrovascular disease, European Society of Radiology, 2011.

Magnetic Resonance Imaging of Perfusion, D. Le Bihan, 1990.

Skull Radiography- Lateral and PA views

2015-01-08T07:42:00.000-08:00

SKULL-LAT VIEW

Patient position

- Patient lies supine to the x ray table.

- Median sagittal plane is parallel to the table.

- The interorbital line is perpendicular to the cassette.

Central ray:

Horizontal ray Centre midway between the glabella and the external occipital

protuberance to a point approximately 5 cm superior to the external auditory meatus.

Essential image characteristics

- The image should contain all of the cranial bones and the first cervical vertebra.

- Both the inner and outer skull tables should be included.

- A true lateral will result in perfect superimposition of the lateral portions of the floors of the anterior cranial fossa and those of the posterior cranial fossa.

- The clinoid processes of the sella turcica should also be superimposed.

1. Coronal Suture

2. Orbital Plates of the Frontal Bone

3. Posterior Clinoid Process

4. Auricle or Pinna (Ear)

5. Lambdoidal Suture

6. External Occipital Protuberance

7. Posterior Arch of C1

8. Sphenoid Sinus

Note: Skull lateral can also be done while the patient is in erect position using vertical ray in the uncooperative patient.

Skull PA (Occipito frontal)

Patient position:

- The patient sits facing the erect Bucky.

- Median sagittal plane is perpendicular to the cassette.

- Neck flex so that the orbito-meatal base line is perpendicular to the Bucky.

Direction of central ray

Centre in the midline of the occiput so that the exit ray emerges in the midline at the level of the glabella.

Essential image characteristics

- All the cranial bones should be included within the image, including the skin margins.

- It is important to ensure that the skull is not rotated which can be assessed by measuring the distance from a point in the midline of the skull to the lateral margin. If this is the same on both sides of the skull, then it is not rotated.

1. Lambdoidal Suture

2. Ethmoid Sinus

3. Petrous Ridge

4. Maxillary Sinus

5. Body of the Mandible

6. Condyle of the Mandible

7. Mastoid Air Cells

With the degree of beam angulation, the position of the petrous ridges within the orbit can be evaluated:

– Occipito-frontal: the petrous ridges should be completely superimposed within the orbit, with their upper borders coincident with the upper third of the orbit.

– OF10° caudally: the petrous ridges appear in the middle third of the orbit.

– OF15° caudally: the petrous ridges appear in the lower third of the orbit.

– OF20° caudally: the petrous ridges appear just below the inferior orbital margin.

Note: Never use skull PA when there is a possibility that the facial bones may be fractured or when the patient is unconscious. When the patient cannot be X-rayed in a prone position, then Skull AP- Supine (frontal-occipital) is used.

Reference:

1. Clark's Positioning Radiography, 12th Edition

Acoustic Noise in MRI

2015-01-04T15:50:00.000-08:00

Acoustic Noise in MRI

-Parajuli K, MSc Medical Imaging, IOM, Nepal

Why noise is a problem in MRI¹

· It causes discomfort, anxiety and distraction to the patient.

· It causes problem in verbal communication.

· It can cause discomfort in sedated patients.

· Some medications can increase the sensitivity of our ear to sound.

· Temporary shift in the threshold of hearing has been reported.

· Neonates have increased sensitivity to acoustic noise.

· Unwanted activation of auditory cortex during functional MRI causes spurious patterns and is a serious problem.²

How is noise produced in MRI

There are many sources of noise in mri.

They can be categorized into four groups

a. Due to gradient current

b. Due to eddy current

c. Due to radio frequency waves and slice selection pulse

d. Due to ambient noise

Noise due to gradient current³

Gradient, spatial variation in magnetic field is produced by the electric currents flowing through coil. Large pulsed electrical current ranging from 200 A to 600A with rise time in the order of sub millisecond or millisecond is applied to copper windings held together by epoxy resin and fiber glass. Gradient coil is itself situated in the magnet bore assembly and thus experience the main magnetic field, a force called Lorentz force develops in between the gradient current and static magnetic field, which causes bending and buckling the coil producing compressional force. Force of 2000 N per meter of gradient coil is produced resulting in vibrations with acceleration levels in the order of 100 m/s². Lorentz force is given by the relation:

where dl is the element inside the magnetic field.

Fig 1: grey arrow showing the direction of gradient current and white arrow shows direction of Lorentz force.

Figure (2) shows two mode of vibration: cone shaped vibration in Z gradient (A) and banana shaped vibration in X-Y gradient (B).

Studies have shown similarity between spectrum of gradient current and that of noise. Also, amplitude of vibration is higher when frequency of gradient current matches the resonant frequency of gradient coil which depends on properties of length and elastic properties of material in coil assembly. It is thought to be due to better acoustic transfer function of the gradient system at resonant frequency. The time course of the gradient current depends on various pulse sequence parameters (TR, TE, FOV). Thus pulse sequence can be modified such that frequency of gradient current (also the frequency of vibration) does not match the resonant frequency of the coil. In EPI the frequency of readout gradient can be adjusted such that it is not same as resonant frequency of the gradient coil.

In addition, the wave for gradient can be changed from trapezoidal to sinusoidal.

Noise due to eddy current⁴

Gradient magnetic field induces eddy current, according to faradays law of electromagnetic induction, in the metallic structures located in the vicinity of the gradient coil for eg. inner cryostat in superconducting systems and Rf body coil, Rf shield etc. These structures conducting the eddy current are situated in a magnetic field because of which Lorenz force develops and induces vibration. Recent study by Eldestein et al ⁴have shown that main source of noise in MRI is due to eddy current induced vibration of the conducting structures located near the gradient coil.

RF Hearing¹

It is due to pressure waves generated because of expansion of tissues as a result of rise in temperature (1x10^-6 degrees C) caused by the RF induced heating. It is very nominal compared to gradient field induced acoustic noise and there is no proof regarding harmful effects due to RF hearing.

Ambient noise³

Inside the scanner room there are various other sources of noise for example:

1. Noise due to air blower of air handling system that can produce noise upto 20 dbA.

2. Pounding sound of low frequency due to liquid helium pump that can reach up to 80 dB.

How much is the noise produced in MRI^{5, 6}

Noise level in MRI depends upon the hardware construction and surrounding environment i.e the presence and absence of various structures that reflect sound. Thus sound level is measured in specific pressure level or sound pressure level which can be expressed in linear or A-weighted scale. The sensitivity of human ear depends upon the frequency of sound. A-weighted sound takes into account the different sensitivity of human ear to various frequencies of sound and is expressed as dBA or dB(A).dB(A) is better correlated with the risk of noise induced hearing loss.

Noise levels of 70 to 130 dB have been reported. Price et al found noise level to be 82.5+-0.1 dB (A) in 0.23 to 0.5T systems and 118.4 +-1.3 dB (A) in 3T systems. Hattori et al⁶ found that noise in 3T MR systems exceeded that of regulatory limit of 99 dB given by IEC. The peak sound pressure level was 125.7 dB for MR angiography, 130.7dB for SS-EPI (single shot spin echo EPI) on the linear scale.

What factor affects level of noise produced?

It depends upon the scan parameters like section thickness, FOV, TR and TE. It is because change in these parameters bring about change in gradient output (rise time and amplitude).Smaller the section thickness and FOV and shorter the TR, higher is the level of noise produced. It increases logarithmically with increase in magnetic field.

Studies have shown presence and absence of patient can also create difference in level of acoustic noise produced. The value was found to be larger in the former case.

What are the regulatory limits?

According to US FDA guidelines peak un-weighted sound pressure level should not be greater than 140 dB. Also, A-weighted root mean square sound pressure level should be greater than 99dBA with hearing protection in place.

UK has provided guidelines for operators as well, according to which hearing protection should be provided to the operators who can be exposed at the average of 85 dB over 8 hour a day.

Exposure limits given by Occupational Safety and Health Administration (OSHA)¹

85dB(A) 16hrs a day

90dB 8 hrs a day

95db 4 hrs a day

100dB 2 hrs a day

105dB 1 hrs a day

110dB 0.5 hrs a day

115dB 0.25 hrs a day

What are the noise reduction strategies?

1. Modification in design of hardware

a) Use of active-passive shielded, vibration isolated and vacuum enclosed gradient coils¹

Gradient assembly can be shielded so that eddy current is not induced in the surrounding metallic structures in the MRI assembly. Active shielding refers to the use of secondary coil other than primary gradient coil(X, Y, Z) that oppose and cancel out the fringe field. Like primary coils there are three shielding coils. Passive shielding refers to the use of conducting layers that are attached on the outer radius of the gradient coil.

Gradient coil are arranged such that structural or mechanical propagation of vibration into other component is reduced this is called vibration isolated gradient coil. Also they can be enclosed in vacuum to prevent air conduction of the vibration.

b) Use of quiet gradient coil

Current in the gradient coil are passed such that net Lorentz force is reduced.

2. Passive noise control

Ear plugs and headphones can be used to lessen the intensity by 10 to 30 dB. Disadvantage of this method is that it interferes the verbal communication with patients.

3. Active noise control⁸

It utilizes the sound, also called anti phase sound that interferes destructively with the sound perceived by the subject undergoing MRI. There are two main topologies for active noise control. Feed forward and feed-back.

Feed-back topology:

· A microphone is placed near to the area where attenuation is required.

· It receives the sound generated by scanner which is perceived by the subject and sends reference signal to the loud speaker.

· Loud speaker, on the basis of reference signal sends control sound that interferes with the noise and produces cancellation effect.

· Disadvantage of this method is that because of the time delay between production of control sound and reception of the same by the subject, high frequency noise cannot be adequately suppressed, making this sort of topology not suitable for EPI sequences where noise of frequency 3 -4 kHz are generated.

Feed forward topology

· Microphone is located near to the sound source which sends the reference signal in advance to the loudspeaker.

· Another microphone called “error” microphone is located near the subject and provides electrical copy of residual noise.

· Advantage of this topology is that higher frequency noise can be attenuated because the reference signal is available on advance and can compensate for the delay due to time required for the propagation of sound from loudspeaker to error microphone. Thus this sort of topology can be used to reduce high frequency noise from EPI sequences.

4. Use of quiet pulse sequences⁹

Pulse sequence with less gradient pulsing can be used. For example Gradient pulsing can be reduced by using single shot sequences based on stimulated echo (STEAM).Further reduction in gradient pulsing can be done by modifying the projection reconstruction method in which the two gradient pulses are replaced by a single mechanically rotating DC gradient coil.

Use of soft gradient pulse is another option. Hennel et.al¹⁰ used band limited pulses that used sinusoidal ramp with least number of slope (ramp) and maximum duration of slope (ramp duration). Noise level was brought to as low as 30dBA in GRE and SE and 60dBA in RARE sequence. Authors named such pulses as soft pulse which can be obtained by convolving the hard pulses i.e the pulses with rectangular waveform with the cosine window of length ᴦ (Figure 3)

Thus there is a growing concern regarding the detrimental effects of acoustic noise generated in the MR systems and the ongoing advancements in MRI hardware, especially in the gradient coil technology, are dedicated towards development of a quiet scanner.

References

1. MC.Jury M,Shellock F.G Auditory noise associated with MR procedure:A Review J Magn Reson Imaging.2000 12:37-45

2. Measuring MRI noise

3. Adriaan Moelker and Peter M.T. Pattynama Acoustic Noise Concerns in Functional Magnetic Resonance ImagingHuman Brain Mapping .2003;20:123–141

4. Edelstein WA, Hedeen RA, Mallozzi RP, El-Hamamsy SA, Ackermann RA, Havens TJ (2002): Making MRI quieter. Magn Reson Imaging 20:155–163.

5. Roozen N.B ,Koevoets A.H and den Hamer A.J Active Vibration Control of Gradient Coils to Reduce Acoustic Noise of MRI Systems.

6. Hattori Y,Fukatsu H,Ishikagi T :Measurement and evaluation of the acoustic noise of a 3 Tesla MR scanner Nagoya J.Med.Sci,2009;69.23-28

7. Tomasi DG¹, Ernst T Echo planar imaging at 4 Tesla with minimum acoustic noise. J Magn Reson Imaging. 2003;18(1):128-30.

8. Chamber J,Bullock D,Kahana Y.Developments in active noise control sound systems for magnetic resonance imaging.Applied Acoustics 2007;68.281-295

9. Alibek S,Vogel M,Sun W etal.Acoustic noise reduction in MRI using silent scan:an initial experience.Diagn Interv radiol 2014;20:360-363

10. F. Hennel,* F. Girard, and T. Loenneker ‘‘Silent’’ MRI With Soft Gradient Pulses Magnetic Resonance in Medicine ,1999;42:6–10

Radio Waves in Magnetic Resonance Imaging (MRI)

2014-10-25T04:53:00.000-07:00

Radio waves in MRI

Author : Ms. Kalpana Parajuli

MSc Medical Imaging Graduate

It is true that we, who are associated with field of medical imaging, are mainly concerned with detrimental effects of ionizing electromagnetic radiation. It seems that our attention is biased and it is high time we realized we are surrounded by non-ionizing radiation and we should know how they interact with biological tissue. This article will mainly focus on radio waves, electromagnetic waves of frequency ranging from 9 kHz and 300 GHz that are used to perturb the longitudinal magnetization in order to produce MR signal. Those of us who belong to this dynamic and ever progressing field of medical imaging should know that the way by which radio wave interacts with the matter is quite different from the pathway that ionizing radiation such as x radiation takes. It is because these waves lack enough energy needed for compton or photoelectric interaction. For the ionization or the breakage of covalent bond to occur a single photon should interact with the electrons in the atomic orbital which is not possible with RF waves that releases its energy through the interaction of multiple photons. The effects of RF energy in human body can be divided into two categories.

a) Thermal also called dielectric heading:

Heating effect of RF waves is best explained by the phenomenon polar molecules show under an electri field. Radio waves are nothing but the oscillating electric and magnetic fields. Under the sinusoidally changing electric field, the polar molecules, molecule with non-zero dipole moment, begin to rotate, as a result of which collision between molecules occurs followed by energy transfer from these molecules to the adjacent molecules. The resultant agitation and energy transfer cause increase in temperature and production of heat respectively. Similarly changing magnetic field induces the electric field (Faradays law) in human body that deposits the energy in the form of heat as described above.

The behavior of electric and magnetic field of radio waves depends on the distance between the source of radiation and the object on which it is incident, as well as on its frequency. The patient in MRI is exposed to waves of frequency that range from 8.5 to 340 MHz and the scanner is in the “near field” i.e “d” is less than one wavelength so heat production is mostly by the magnetic field of RF waves with little or minimal effect of electric field of RF wave. The reason behind this sort of dominance is the independent relationship between “E” and “B” in the near field.

b) Non-Thermal :

These are specific effects that occur due to interaction of magnetic and electric field vector with human body other than heating. Very less is known about non-thermal effects because of lack of conclusive scientific evidence in human models and therefore not taken into account by regulatory bodies while providing guidelines on safety limit.

Specific Absorption Rate (SAR) is the name of dosimetric quantity used to describe the rate of heat deposition in unit mass of tissue, and is measured in W/kg.

What affects SAR in MR imaging?

The answer to this question is given by the figure below.

It is thus clear from above formula that SAR is the function of patient related factors as well as scan parameters. SAR increases with the square of magnetic field strength of the magnet system of scanner, because high power radio waves are needed to cause resonating effect on the spins of higher larmor frequency. It also increases with the square of the flip angle, whereas it varies proportionally with the duty cycle (ratio of average to peak RF power) and patient weight. Because conductivity and tissue density vary from person to person, the calculation of SAR is not easy and accurate. Animal experiments have shown that the permittivity and conductivity of tissue decreases with age thus young ones are more vulnerable to deleterious effects of radio waves. It also depends on the perfusion status and geometric configuration of exposed tissue and presence of metallic implants. Dependency of SAR on several factors is one reason behind the challenges in RF dosimetry.

SAR of 1W/kg is said to cause increase in temperature of insulated object (phantom) by 1⁰C in an hour. In human and animals, input of 4W/kg of SAR has shown to raise the temperature by 1⁰C. It is not practical to measure change in temperature (core/whole body or localized) thus SAR is used to quantify the RF exposure. International Electrotechnical Commission( IEC 60601-2-33) and Food And Drug Administration (FDA) has given limits for both temperature as well as values of SAR (shown below). The guidelines consist of limits for the whole body and local level exposure, because some organs are highly heat sensitive than others because of higher resistance, and are less affected by thermoregulatory system of body because of less perfusion, for example: eye, gonads, and thus separate limit for local RF exposure was required.

It is now realized that separate guideline is necessary to account the change (increase) in SAR by metal implanted in the body of patients. However none of the current guidelines have addressed this issue. The limits are also given for occupational exposure. Those given by (IEC) are same as that for patients whereas that by Institute of Electrical and Electronics Engineers (IEEE) and International Commission on Non-Ionizing Radiation (ICNIRP) are one tenth of the maximum limit for patients.

Just like CTDI volume and Dose Length Product are displayed in the CT scanner, SAR value is also displayed on the monitor of the MRI system. There are many methods available for estimation of SAR. Two basic methods are caloriemetric method and pulse energy method. At the beginning of the scanning the machine runs calibration to find out the energy required to flip the spins by 90 and 180 degree. Power is then obtained by dividing the total energy of all pulse in one sequence with time of repetition (TR), and the result is ultimately divided by weight of the patient to get SAR. This is why the scanner requires us to input the value of weight of the patient.

Till now we believe that increase in patient weight increases the SAR and all of us have habit of looking at the value of SAR given by the machine itself to determine whether our protocols are safe. However results of a recent study were quite astounding because negative correlation was observed between patient weight and the SAR calculated by 3T scanner, whereas in 1.5 T scanners the relationship was maintained. This study has indeed raised a big question mark on the reliability of the values of SAR provided by the manufacturers of MRI scanner.

What about the consequences of RF induced heating in human?

Many incidents of second and higher degree burn in patients undergoing examination in 1.5 and 3.0T MR systems have been reported. While documentation of adverse consequence of RF associated heating of the metallic implants are available, other physiologic response for example- changes in heart rate, oxygen saturation, blood pressure, respiratory rate and cutaneous blood flow has not shown to cause effects that needs serious concern.

What are the possible ways to reduce SAR in MRI?

We are familiar with the tradeoffs between the radiation dose in CT and the image quality. A balance between them is essential part of protocol selection. Similarly in MRI as well, SAR can be minimized through wise selection of parameters, which however also affects other areas like imaging time, image quality etc. For example: reducing the flip angle affects the image contrast, reducing the no of slices increases imaging time. Other methods can be; reducing the echo train length of turbo spin echo or fast spin echoes, the use of quadrature rather than linear coils for

transmission. Parallel imaging technique, by using multiple receivers increases the amount of data received and thus reduces the imaging time by a certain factor (reduction factor) without the use of additional RF pulse and thus reduces SAR.

In conclusion safety issues regarding use of radio waves in MRI are gaining much more attention than ever mainly because of three reasons - 3T scanners, that needs high power RF amplifiers, about 35 kW, are widely being used; the popularity of new RF intensive sequences (HASTE, FIESTA, true FISP) have also heightened and MRI is no longer contraindicated for the patients with implanted materials. Thus it is necessary that manufacturer, the scientific research community and user should collaborate with each other to prevent the potential hazardous situation that can arise due to RF energy in MR environment.

MSc Medical Imaging Graduate Ms. Kalpana Parajuli

Maharajgunj Medical Campus (MMC),

Tribhuvan University Teaching Hospital (TUTH),

Kathmandu, Nepal

What is radiation quality?

2014-10-10T03:06:00.000-07:00

What is radiation quality?

A measure of the damage potential of a particular radiation type,this means alpha betas and gammas and positrons for example,has different abilities to cause damage to tissue,this damage potential is related to its size, charge and energy,we have a good example of this with amunitions, let's take a unit that you already know well,

energy gained by electron being accelerated through a field of 1 V,now to begin the radiation damage to tissue, let's first see the linear energy transfer,it is a measure of energy absorbed by the material, keV/length...somewhat related to the track or path.let's consider a beta particle moving through a particle such as lead brick,it should be intuitive and it interacts, it transfers it's energy to the material.high LET=more energy transfer and shorter distance travelled.Because it loses energy quickly.ability to produce more damage,of course we are concerned with biological damage,before we begin to understand the damage of radiation, we must discuss radiation dose measurements in radiation symmetry.

radioactivity rate of decay= number of disintegration per unit time (usually sec)
exposure=the electric charge in air due to the presence of radiation, the total sum of kinetic energy of charged particles liberated by uncharged radiation.

absorbed dose=similar to exposure-consider tissue type
energy deposited in tissue.value of absorbed dose from same radiation will change with tissue type.

1Roentgen = 1Rad
100rad=1Gy
in human tissue, about 87% of 1R is absorbed 0.87R=1rad
dose equivalent=corrects absorbed dose for damage potential of radiation type
must provide a weight factor of radiation types.
QF or RBE,gives numerical value to the damage potential.
QF=dose from standard radiation to produce given biological effect/dose from test radiation to produce same biological effect
standard radiation -200 kVp x-rays
dose equivalent:
rem \ sievert
100 rem=1Sv,
Dose equivalent =absorbed dose x QF or RBE
ICRP defined values for QF/RBE
gamma x rays 1
electrons (beta/positron) 1
neutrons 10
alpha 20
QF INCREASES as LET increases
150mrad x 1 gamma - 150 rem
75 mrad x 20 alpha = 1500 rem
total dose equivalent = 150 + 1500 rem
for QF =1
general rule 1 R =1 rad = 1 rem
reality:
1 rad = 0.96 rem or 0.87 R = 1 rad
there is no universally accepted conversion from rad to rem.
exposure: air
absorbed dose: considers material
dose equivalent: considers radiation type
effective dose equivalent:
measures the biological risk on the whole body by considering different tissue types response differently to radiation ; also accounts for radiation to only part of the body vs the whole body.

to account for different tissue response, weight factors are used (yes, again), weight factors determined by ICRP.
2007 revision of ICRP recommendations:
14 designated tissues for organs + 1 for everything else.
to calculate: summation of- Dose Equivalent x Weight Factor
For all designated tissues/organs irradiated
H = summation Wt x Ht
How do we account for long term dose to an organ? inhaled, injected or swallowed
committed dose equivalent: Ht,50, the dose equivalent to an organ or tissue received during the 50 year period following intake of the radioactive material, we use 50 because there isnot so much damage after 50,as the receipent would likely to be perished by then.committed dose equivalent accounts for physical decay and body excretion.

biological half life: SAME principle as radioactive or physical half life,and it is how we account for body excretion of radioactive material,it is important to know that it is an approximation differs from person to person,sickness are different and medications are different,

Biological Half life:depends on the renal system and the bowels, the overall disappearance of radioactivity from the body depends on effective half life, consider both physical and biological decays,
1/effective halflife=1/biological halflife + 1/physical halflife
we have seen many ways to calculate dose now considering many variables.
it becomes very complex trying to calculate a specific dose to a specific organ.
source organ is lung, target organ is liver.after.....then same organ liver can be source and target.
to calculate dose to the liver precisely we need to answer few questions//
exact location in lungs, time spent in lungs, exact location lin liver, time spent in liver.
doses from each adjacent organ of liver must be considered, which becomes even more complicated.
another factor we must consider in calculating dose to tissue is the percentage of radioactive emissions that deposit energy in the tissue vs those that escape.this value is called the absorbed fraction.
depends on track, density and thickness, radiation type, energy and half life.
absorbed fraction will be expressed as a number b/w 0-1, where 1.0 equals 100% absorption.
with all of these factors, how can we ever calculate an accurate dose to an organ?
MIRD - Medical Internal Radiation Dose Committee:which is the permanent committe of nuclear medicine provided a widely accepted method to calculate internal dose,
the first method is to calculate cumulated activity;plot activity vs time, area under the curve is the cumulated activity;units = activity x time uCi x hr
it is important to understand that the cumulated activity is the first step in determining the dose to an organ that a patient receives from a nuclear medicine procedure.
S factor: consider most of the physical aspects of location of radioactive material,
total energy associated with the radiation typexabsorbed fraction in target of source/mass of the target organ.
complications of calculating S factor: absorbed fraction varies greatly with gamma rays, also all transitions, disintegrations and emissions must be included not just the primary decay, for 99mTc, there are altogether 14 radiations to include.this include conversion electrons, fluorescent x-rays and auger electrons,

lucky for us, MIRD published all the values of S-factor, so we don't need to calculate all these values.
DOSE equation of MiRD=cumulated activity x S-factor
Cumulated Activity = Biological Parameters
S-factor = Physical parameters
review:
LET and its affect on dose
all measurements of radioactive energy deposit
unit conversion b/w SI and traditional
how to incorporate QF in dodse calculations
how to calculate effective half life
how to calculate organ dose using MIRD scheme

Reference:

https://www.youtube.com/watch?v=MTPTmIopuA0

Geiger Muller - Gas Filled Detector

2014-10-05T09:47:00.001-07:00

Geiger Muller - Gas Filled Detector

Geiger Muller or counters are one of the kind of gas filled detector;that operate by using the ionising nature of alpha, beta and gamma radiation;neutron sensitive devices can also be produced;typically by introducing boron;which interacts with neutrons to produce secondary ionization particles that trigger the counter response. The GM tube is a shield metal tube cylinder with low pressure inert gas such as Argon or Neon;a thin metal wire runs down the center of the tube; which is electrically insulated from outer cylinder till the end of the tube; the front of the tube is shield with radiation window,that is specific to typical radiation that is to be detected by the counter;a thin micro window is to be used if the tube is to be sensitive to alpha particles and low energy beta particles;both of which have low penetrating power;a thicker window of different material such as glass or thin sheet of metal is to be used to measure high energy beta particles while for gamma rays.

The tube is often shield without a window;in such tubes, the detection occurs when high energy photons liberate electrons from the tube outerwall,the inner wire and outer wall are maintained at a pd of about 1kV and in the absence of radiation, no current can flow through inert gas b/w the central anode and the outer cathode; the connections are made by wires into a connection housing;fixed over the rare of the tube;and the outer tube guard is actually screwed onto the tube to protect the actual GM tube;this tube guard can be open at the end or can be covered with an energy filter;to change the energy and particle sensitivity of the device or a careful calibration is used, it allows for ambient dose measurement;rather than ambient count measurement to be made;the wires are connected to the tube to the control electronics;by power to perform counting operations and allow functions such as conversion from counts to dose, data logging, data averaging and driving the display;the tube works on the basis of gas amplification,incoming radiation ionizes some of the inert detector gas resulting in a free electron and a positively charged ion; the electric field inside the tube attract the charged ions towards outer cathode while the electrons are attracted towards central anode, as the electrons encounter the anode, the electrons experience the growing strength so that their accelerating forces increases, near the anode the acceleration is such that enough energy that it either excites the electrons in other atoms of the detector gas,or ionize them completely, excited electrons quickly decay releasing photons that can trigger ionization farther along the tube while the electrons free due to ionization can go on to cause further ionization leading to an exponential growth; this is often referred to as avalanche effect.

The charge migration in the tube leads to reduction of potential in an anode and increase in potential of the cathode either of which may be detected as a signal by the counter electronics,as the negative charge around the anode increases, the effective electric field is reduced, eventually this reduction is such that further avalanche are not possible, and the tube can no longer detect radiation,this step persist until the sufficient electrons have recombined with the anode, and the positive gas ions recombine with the cathode so that the field is recoverable enough in strength to trigger another avalanche,this is so k/a dead time of the detector,the time after detection the counter is insensitive to further events and this existence means that the detector count rates must be corrected to give the actual count rate,after the dead time further detection are possible with reduced signal strength, the total time that elapses before the full strength signal is produced by subsequent events is k/a recovery time.

Part 2:
The recombination of positive gas ions at cathode may be problematic as the ions may be neutralized in excited state or dislodge electrons at the cathode,when in excited state the atoms will eventually decay to ground state by emitting a photon and these photons and the dislodged electrons may be the cause of reionisation of the gas triggering another avalanche so the single detection event will lead to a continuous discharge, to prevent this a quenching mechanism is used, the quenching may be electronic so the electric field is removed fro a short period of time following an event to prevent further discharge or may be inherent in the design by mixing quenching gases with the detection gases such quenching gases are easier to ionize than detection gas so that during migration to the cathode, the detection gas is neutralized by the quenching gas which then becomes the positive ion, migrating to the cathode, when the charged quenching gas ions recombines with the cathode, it does so in the ground state so that further avalanche discharge is avoided.

The most effective quenching gas are the organic compounds,but these are dissociated irreversibly during quenching which gives the tube operating limiting longevity, an alternative is to use halogen gas which is recovered full at cathode, so avoid the removal of quenching gas, the raw output from GM tube shows a fraction of radiation counts per sec which is modified by taking in account of the dead time to give actual counts per sec; if the radiation type and energy are known, then the counts per sec can be calibrated so that the unique gives an equivalent dose rate,this is not the best method for dose determination from an unknown source as in the GM tube signal pool side is relatively insensitive to the instant radiation energy type, so the energy deposited is difficult to determine,the use of counts per sec or dose rate depends on large extent on circumstances of use, on both modes the radiation are being detected, in the former the activity is displayed while in the later, conversion is made to indicate the energy deposition rate,

For wide spread contamination by a radioactive material the energy rich in the GM tube will be small as the inverse square law and transient absorption remove all of the fraction of the instant particles however radiation from contamination can be measurable as the increase in the background radiation level which can easily be expressed in the increase in the no of counts per sec, as a practical radiation detection device,GM tube is not considered a natural choice for measuring the radiation than pulse device such as linac, the reason is the dead time is much longer than the pulse width duration, the detector will only pick up a single event rather than a bunch of electrons or photons and the tube will count the pulses not the radiation leading to an underestimation of dose,if that time is significantly longer than bunch frequency, the detector will count only a fraction of bunches,leading to further underreading,if the operating characteristic of GM tube and linac are known, these effects can be compensated for so the GM tube could be used, this will mean however,different methods have been used to measure the different calibration methods to measure the data,

Reference:
https://www.youtube.com/watch?v=bcjMOr-qiwA
https://www.youtube.com/watch?v=1qRjSLqM4zg

Gadolinium : An MRI Contrast Agent

2014-06-29T02:01:00.000-07:00

Gadolinium: An MR Contrast Agent

MRI is based on the response of proton spin in the presence of an external magnetic field when triggered with a radio frequency (RF) pulse. Under the influence of an external magnetic field, protons align in one direction. On application of the RF pulse, aligned protons are perturbed and subsequently relax to their original state. There are two independent relaxation processes: longitudinal (T1) and transverse (T2) relaxation, which are typically used to generate the MR images.¹

In order for an excited spin system to return to its equilibrium magnetization, energy must be transferred from the spin system to the lattice (surrounding). The return to equilibrium is described by the spin-lattice relaxation time (T1). When T1 weighted sequences are used, the magnitude of the MR signal increases with decreasing T1 relaxation times. Further, the contrast between two tissues will of course also increase with increasing difference in T1 relaxation times between the two tissues. We know MRI creates images capable of differentiating among different tissues based on their T1 and T2 properties. However, the inherent difference in T1 relaxation time between biological tissues, or between normal and pathologic tissue is not always large enough to obtain a detectable contrast in the MR image. Sufficient contrast is of particular importance in differentiating pathological tissue from normal surrounding tissue. Exogenous MR contrast agents were therefore developed shortly after the first commercial MR systems became available in the early 1980’s. Today, MR contrast agents are typically in a significant proportion of MR examinations; with the highest usage in CNS applications (tumor diagnosis). MR contrast agents are widely used in MR angiography (MRA), they are injected into the bloodstream, and strongly T1 weighted images are acquired. Blood vessels appear much brighter than any other tissue that highlights the vessels.³

All biological systems are composed of various molecules and organisms which have different proton concentrations and different T1 relaxation times. The presence of paramagnetic ions (e.g., Gd⁺³ and Mn⁺²) near the tissue enhances its relaxation and shortens the T1 relaxation time. Contrast agents with T1 weighted enhancing ability produce bright positive signal intensity in images and increase the conspicuousness of cells, facilitating easy tracking of cells in low signal tissues. Among, those paramagnetic ions, Gd⁺³ are the most effective T1 weighted contrast agent for clinical use.¹

Paramagnetic contrast agents, which have unpaired electrons in the outer electron shell and thus also a magnetic momentum, are used to improve imaging. Electrons affect the protons since they have a 657 times stronger magnetic momentum than protons. As a result, both the proton density in the tissue is changed directly and the local magnetic field is changed indirectly by the interaction between the electron spin of the contrast agent and the surrounding hydrogen nuclei. Apart from the paramagnetic elements manganese and iron, the lanthanide gadolinium, with its seven unpaired electrons in the outer electron shell, interact with protons in nearby water molecules to dramatically shorten the T1 relaxation time and is one of the metals most commonly used in MRI contrast agents. The appearance of tissue in which the contrast agent penetrates appears brighter on T1 weighted imaging relative to non-contrast enhanced tissues. Because of its intrinsic toxicity, however, gadolinium cannot be used in the free ionic (Gd⁺³), but only in the form of its water soluble chelate complexes. In particular, the derivatives of diethylenetriamine pentacetic acid (DTPA) have proved useful as MRI contrast enhancing agents.⁴

MR contrast agents act by selectively reducing T1 (and T2) relaxation times of tissue water through spin-interaction between electron spins of the metal containing contrast agent and water protons in tissue. However, here, we are only discussing about Gadolinium based contrast agents. In the presence of the contrast agent, the observed relaxation rate (R1=1/T1) can be split up in an intrinsic tissue contribution and a contribution from the contrast agent, according to

R1^obs = R1^tissue + R1^ca,

Where R1^tissue is the intrinsic relaxation of the tissue without the contrast agent and R1^cais a paramagnetic contribution of the contrast agent. The contribution of the contrast agent can be written as

R1^ca = r1. (CA),

In which r1 is the relaxivity (in mM^-1s^-1) and (CA) the concentration of the contrast agent. Using R1 =1/T1, this leads to the well-known equation,

which shows that the shortening of the relaxation rate R1 is linear with contrast agent concentration and that contrast in the MR images can be enhanced either by using a contrast agent with a high relaxivity r1 and/or by increasing the local contrast agent concentration. It should be noted that the relaxivity apart from being a contrast agent specific parameter, also depends on the solvent and distribution, which could vary considerably in vivo, e.g. when the contrast agent is confined to the blood pool or compartmentalized in the cytoplasm of cells. This means that the contrast agent may not affect all water protons in the tissue equally. Linearity of R1 with concentration can therefore not be guaranteed under all circumstances.²

The paramagnetic contribution from the contrast agent is generally understood to originate from relaxation in two pools of water coordinated either directly with the Gd⁺³, called the inner sphere, or located in the second coordination sphere and the bulk, referred to as the outer sphere. The detailed interactions between the ion and the water can be understood in terms of the Solomon-Bloembergen-Morgen (SBM) theory.²

At higher concentrations, signal saturation occur meaning that a further reduction in T1 does not result in a further increase in signal intensity since the longitudinal magnetization is fully recovered. Further, when the concentration becomes very high (depending on TE) the signal will start to fall with further increase in concentration since the T2 effects of the contrast agent will start to dominate the signal behavior.⁵

Note that the contrast agent does not leak out from the intravascular space in normal brain tissue due to the presence of a blood brain barrier (BBB) which prevents even small molecular weight molecules like Gd chelates to enter the interstitial. Since the intravascular volume in the brain is small (<5%), little enhancement is thus seen in healthy brain tissue. In brain tumors however, the BBB can often be disrupted due to various pathological processes, resulting in a selective accumulation in the extravascular space in the tumor. The fact that many pathological processes alter the permeability of the BBB, resulting in selective accumulation of MR contrast agents is, in fact, the main reason why these agents are so useful for CNS imaging since the absence/presence of CA leakage as well as the pattern of contrast enhancement can give important indications as to the type of pathology present. Therefore, MR contrast agents not only increase the sensitivity (the ability to detect) but also the specificity (the ability to differentiate) of the diagnostic procedure.⁵

In most organs it passes from the vasculature into the interstitial space relatively quickly. After the initial redistribution into the extracellular fluid space with a half-life of about 11 min, Gd is gradually excreted via kidneys with a biological half-life of approximately 90 min, so in most patients it is not detectable in tissues after about 6 h although it may linger in the urine and bladder for a day. At low concentrations such as those used in normal clinical practice the major effect is the T1 shortening, and tissues which take up the agent have enhanced signal intensity on T1WI. Most clinical sites use a standard dose of 10 or 15 ml for adult patients, who approximates to 0.1mmol Gd per kg body wt. Double and even triple dose injections are routinely used for MRA and perfusion imaging and have been shown to improve the conspicuity of lesions in multiple sclerosis and metastatic disease. At concentrations higher than about 1mmol Gd Kg^-1, ten times the standard dose, the effect on T2 begins to dominate and a loss of signal occurs.⁶

Fig: Signal Intensity vs concentration of Gd, calculated using T1W SEQ (TR=400ms, TE=15ms) and a tissue with T1=800ms and T2=75ms

Magnetic resonance is one of the leading diagnostic imaging modalities, since it excels in depicting tissues with high spatial resolution and has superior soft tissue contrast. Nevertheless, MRI can considerably gain from the use of contrast agents. Contrast enhanced MR angiography is now routinely used in the non-invasive evaluation of vascular diseases. Dynamic contrast enhanced MRI receives considerable attention in the assessment of stroke and tumor vascularity. MRI applications are becoming more and more dependent on contrast agents. Traditionally MR contrast agents are based on their ability to lower T1 and T2 of water. Nowadays, nanoparticle based contrast seems to be a powerful technology in both basic science as well as clinical settings. Chemical Exchange Saturation Transfer (CEST) agents, based on magnetization transfer methods have also been developed where contrast can be switched on and off at will, which facilitates localization and detection of contrast agents in living organism without the need for exact co-registration of pre and post contrast MR images. The sensitivity of CEST has increased with the invention of (liposome based) LIPOCEST agents. Researches are going on to study subcellular processes from bimodal imaging combining fluorescent markers, such as quantum dots, with MR contrast agents.^1,2,5

Reference:

1. Zhu et al, Nanoparticle-based systems for T1-weighted Magnetic Resonance Imaging Contrast Agents, Int.J.Mol.Sci.2013.

2. Strijkers et al, MRI Contrast Agents: Current Status and Future Perspectives, Anti-Cancer Agents in Medicinal Chemistry, 2007.

3. Image Contrast, International Center for Postgraduate Medical Education, 2009.

4. Peter et al, Gadolinium based contrast agents, Metrohm International Headquaters, Herisau, Switzerland.

5. MR Contrast Agents, FYS-KJM 4740 – The Physics of MRI.

6. McRobbie et al, MRI from picture to proton, Cambridge University Press, Second Edition, 2006.

Basic Principle Of CT Scan

2014-06-28T22:18:00.001-07:00

Magnetic Resonance Coils (MR Coils)

2014-06-24T19:45:00.001-07:00

Basic Principle of Mammography (Physics Concept)

2014-06-16T01:11:00.001-07:00

CT Scan Instrumentation

2014-06-12T21:15:00.001-07:00

CT SCAN GENERATIONS with latest advancement

2014-06-12T20:43:00.001-07:00

Intravenous Urography (IVU) Techniques

2014-06-12T20:02:00.001-07:00

Quality Control in Digital Radiography and Mammography Part 1

2014-04-23T02:53:00.001-07:00

Radiography of Upper Limb

2014-04-22T00:15:00.001-07:00

Barium Enema and Loopogram

2014-04-20T08:26:00.001-07:00

Principle of Ultrasound Physics

2014-04-18T22:56:00.003-07:00

Principle of Ultrasound Physics

2014-04-18T22:56:00.001-07:00

CT Procedure of lower limb

2014-04-17T13:29:00.001-07:00

Principle of SPECT and PET Physics

2014-04-17T07:55:00.001-07:00

Principle of CT scanning Physics

2014-04-16T08:31:00.001-07:00

Barium Follow Through and Small Bowel Enema

2014-04-07T19:46:00.001-07:00

Image Quality in x-ray film

2014-04-06T20:07:00.001-07:00

Picture Archiving and Communications

2014-04-03T07:53:00.001-07:00

X ray artifacts

2014-04-03T05:09:00.001-07:00