Medical x ray image sensors

Medical X-Ray Imaging X-Ray Digital Image Sensors

Presented by: Islam Kotb

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Presentation Agenda

1. X-rays: A brief history 2. How X-ray works 3. Types of medical X-rays 4. Image Detection 5. Flat Panel Detector

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1. X-rays: A brief history

Over one hundred years ago on November 8 1895, the German physicist Wilhelm Conrad Roentgen (figure 1.1)happened upon X rays in his laboratory in Würzburg.

On December 28 1895, Roentgen announces his discovery with a scientific paper, W. C. Roentgen: About A New Kind of Rays.

1896 Fluoroscopy is invented in January by Italian scientist Enrico Salvioni, while American inventor Thomas Edison works on a similar device

Figure 1.1:Wilhelm Conrad Roentgen

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Figure 1.2:The famous radiograph made by Roentgen on December 22, 1895. This is traditionally known as “the first X-ray picture” and “the radiograph of Mrs. Roentgen’s hand.

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1901 Roentgen wins the first Nobel Laureate in Physics prize for his discovery.

1904 Clarence Dally, Thomas Edison’s assistant in X-ray research, dies of extreme and repeated X-ray exposure

1910 Eye goggles and metal shields are commonly used to shield X-ray users

1919 Dr. Carlos Heuser, an Argentine radiologist, is the first to use a contrast medium in a living human circulatory system.

1920–1929 Chest X rays are used to screen for tuberculosis.

Roentgen dies February 10, 1923. 1927 Portuguese physician, Dr. Egaz Moniz the first to

create images of the circulatory system in the living brain.

Drs. Evarts Graham and Warren H. Cole of Washington University,St. Louis, discover in 1923 how to visualize the gall bladder with X rays by using contrast media.

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In 1936, the first “tomograph”—an X-ray “slice” of the body—is presented at a radiology meeting.

The Betatron, a circular electron accelerator, is developed by Dr. Donald Kerst of the University of Illinois between 1940–1943.

In 1960, Dr. Robert Egan of the University of Texas M.D.Anderson Tumor Institute, Houston, with the support of the U.S. Public Health Service, publishes the results of an intensive, three-year study of mammography.

1970–1979 CT, or computed tomography, which takes X-ray “slices” of the body and images them on a computer screen, is introduced.

1980s MRI (magnetic resonance imaging; also referred to as MR)—the marriage of a strong magnet and a computer—is introduced, instead of X-ray’s ionizing radiation

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1980s the development of digital technologies for X-ray imaging

1995 The introduction of Flat Panel Detector (FPD)

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Presentation Agenda


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2- How X-ray works Conventional x-ray radiography produces

images of anatomy that are shadowgrams based on x-ray absorption.

The x-rays are produced in a region that is nearly a point source and then are directed on the anatomy to be imaged.

In medical x-ray imaging, the x-ray energy typically lies between 5 and 150 keV, with the energy adjusted to the anatomic thickness and the type of study being performed.

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Figure 2.1 X-ray tube

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The x-rays emerging from the anatomy are detected to form a two-dimensional image, where each point in the image has a brightness related to the intensity of the x-rays at that point.

Image production relies on the fact that significant numbers of x-rays penetrate through the anatomy and that different parts of the anatomy absorb different amounts of x-rays.

In cases where the anatomy of interest does not absorb x-rays differently from surrounding regions, contrast may be increased by introducing strong x-ray absorbers. For example, barium is often used to image the gastrointestinal tract.

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Figure 2.2: Overview of Radiography process

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Presentation Agenda


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3-Types of medical X-rays

3.1 Analog Radiology

Figure 3.1 Analog X-ray

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3.2 Digital Radiology

Figure 3.2 Digital X-ray

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3.3 Mamography

Figure 3.2 Digital X-ray

Figure 3.3Schematic diagram of a dedicated mammography machine.

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3.4 Computed Tomograpgy (CT) In the mid 1970S, CT gave physicians a whole

new way of seeing. By eliminating the interfering patterns that

come from over- and underlying bones and organs, CT provides ample contrast among the various soft tissues.

CT is routinely used for detailed studies of abdominal and pelvic organs, the lungs, and the brain.

CT can image objects down to about 1/3 millimeter

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Figure 3.4 Computed tomography. This view of a slice of bone, eyes, and brain tissues several millimeters thick displays good soft-tissue contrast and detail, and low visual noise.

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Figure 3.5 CT scanner machine

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3.5 Flouroscopy Fluoroscopy is radiography's first cousin. Here, the X rays that pass through and emerge

from the patient are projected onto the front face of an image intensifier, an electronic vacuum-tube device that transforms a life-size pattern of X-ray shadows into a small, bright optical image.

This visible image can be fed into a film camera; more commonly, it goes to a television (video) camera, where it is converted into an electrical signal and sent to a video monitor for live display.

The image can be recorded on videotape for subsequent playback and further processing.

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Figure 3.6 Flouroscopy

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Figure 3.7 Different types of Flouroscopy machines

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3.6 Dental X-rays

Figure 3.8 Types of dental x-ray

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Presentation Agenda


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4- Image Detection 4.1 Screen Film 4.2 X-Ray Image Intensifiers with

Televisions An x-ray image intensifier detects the x-

ray image and converts it to a small, bright image of visible light.

This visible image is then transferred by lenses to a television camera for final display on a monitor.

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Figure 4.1 X-ray image intensifier

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4.3 Photostimulable phosphors (PSPs) They were pioneered by Fuji in the 1980s. In modern hospitals a photostimulable phosphor

plate (PSP plate) is used in place of the photographic plate.

After the plate is X-rayed, excited electrons in the phosphor material remain "trapped" in "colour centres" in the crystal lattice until stimulated by a laser beam passed over the plate surface.

The light given off during laser stimulation is collected by a photomultiplier tube and the resulting signal is converted into a digital image by computer technology, which gives this process its common name, computed radiography.

The PSP plate can be reused, and existing X-ray equipment requires no modification to use them.

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Figure 4.2 PSPs and the reader

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4.4 Flat Panel Detector (FPD)

The flat electronic detector provides direct digital registration of X-ray images, without the intermediate stage of optical or mechanical scanning.

It has the advantages of : 1. It has a high DQE, allowing low-dose operation 2. It has a compact construction for easy integration 3. It offers a new perspective for live X-ray

examination 4. High image quality 5. The integrated flat a-Si detector provides a

significant increase in efficiency, particularly in busy bucky rooms.

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Presentation Agenda


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5-Flat Panel Detectors (FPDs) FPDs are classified in two main categories: 1. Direct FPD's 2. InDirect FPD's In all the methods, the charge is accumulated

for a frame period before being read out. Gamma cameras, in contrast, count each x-ray

photon as it arrives. That technique is generally not used for x-ray

imaging because the x-ray photon arrival rates are too high to permit counting

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5.1 Direct FPDs: There are 2 types of direct FPDs: 1. The Intrinsic Method

Figure 5.1: Intrensic Direct FPDs

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Arriving x-rays are captured by the amorphous silicon diode where hole-electron pairs are generated.

An applied bias separates the charge to prevent recombination.

Because a charge pair is generated for about each 5 electron volts of x-ray energy, the signals are high.

Unfortunately, the x-ray absorption of silicon is very low so the photodiode needs to be 10 to 20 mm thick.

Fabricating such devices of amorphous silicon is not feasible.

Intrinsic devices have been made from crystalline silicon but only arrays of one or two lines are practical and even these are expensive.

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2. The Photoconductor Method

Figure 5.2: a-Se Direct FPDs

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Photoconductive materials with higher x-ray absorption than silicon can be coated on an array of conductive charge collection plates each supplied with a storage capacitor.

These also produce hole-electron pairs when x-rays are absorbed but the charge generated must be stored out of the layer to avoid lateral crosstalk.

The applied field not only separates the charge but directs it towards the collector plate directly below to maintain image sharpness.

Currently, the only photoconductor in production, selenium also has relatively low x-ray absorption and requires about 50 electron volts to produce a hole-electron pair.

These restrict both the minimum dose needed and the size of the signal generated.

Other materials with lower energy requirements and higher x-ray absorption are under development.

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5.2 Indirect FPDs:

Figure 5.3: Indirect Flat panel detector signal chain.

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This method incorporates the usage of a scintillator.

A scintillator is a compound that absorbs x-rays and converts the energy to visible light.

A good scintillator yields many light photons for each incoming x-ray photon; 20 to 50 visible photons out per 1kV of incoming x-ray energy are typical.

Scintillators usually consist of a high-atomic number material, which has high x-ray absorption, and a low-concentration activator that provides direct band transitions to facilitate visible photon emission.

Scintillators may be granular like phosphors or crystalline like cesium iodide.

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The FPD consists of a sheet of glass covered with a thin layer of silicon that is in an amorphous, or disordered state.

At a microscopic scale, the silicon has been imprinted with millions of transistors arranged in a highly ordered array, like the grid on a sheet of graph paper.

Each of these TFTs is attached to a light-absorbing photodiode making up an individual pixel (picture element).

Photons striking the photodiode are converted into electron-hole pairs.).

Since the number of charge carriers produced will vary with the intensity of incoming light photons, an electrical pattern is created that can be swiftly read and interpreted by a computer to produce a digital image.

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5.3 Scintillators

Figure 5.4:The relative light spreading of columnar CsI versus an amorphous phosphor screen.

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Phosphors are materials which glow when exposed to x-rays.

For maximum brightness, the phosphors use in x-ray imaging are made of rare-earth oxysulfides doped with other rare earths.

The most common are gadolinium and lanthanum oxysulfides doped with terbium.

These typically emit blue to green light which is well-matched to film sensitivity.

Various grain sizes and chemical mixtures are used to produce a variety of resolution and brightness varieties. In use, these are mixed with a glue binder and coated on to plastic sheets.

Tens of electron volts are needed to produce each visible photon in a phosphor screen and x-ray absorption is good.

Light scatter can be a problem if the layers must be thick to stop higher-energy x-rays.

A.Structure of a phosphor scintillator:

Figure 5.5: Phosphor scintillators

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For a better combination of resolution and brightness, cesium iodide is used.

CsI grows as a dense array of fine needles under the proper evaporation conditions.

This produces crystals which act as light pipes for the visible photons generated near the input side of the layer allowing very thick layers to be used with excellent retention of resolution.

About 20-25 electron volts are needed to generate each light photon.

When doped with thallium, CsI emits at about 550 nm, just at the peak of the spectral sensitivity of amorphous silicon.

The combination of CsI and amorphous silicon has the highest DQE of all materials in production today.

B.Structure of CsI scintillator:

Figure 5.6: CsI scintillators

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5.4 Architecture of Indirect FPD

Figure 5.7: FPD internal architecture.

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Closely coupled to the array is the X-ray scintillator. Generally, rare earth screens such as gadolinium

oxysulfide, can be a separate detachable sheet which is mechanically forced into close contact with the array.

If a CsI screen is used, this is often directly deposited on the array, to give the best optical coupling efficiency.

CsI is used in applications like low-dose fluoroscopy, where the photon flux is very low.

Figure 5.8 shows a comparison between the absorption efficiency of CsI and gadolinium oxysulfide.

In addition to its much higher absorption efficiency, CsI also produces roughly twice the light output of a gadolinium screen, which results in more than four times the signal at the photodiode for a given dose.

5.4.1 Scintillator layer

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Figure 5.8:Absorption efficiency of the primary scintillators used in FPDs

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The thickness of the CsI can be greater than that of a rare earth screen because when CsI is deposited on the array it grows in a columnar structure.

The columns tend to act as light pipes, reducing the amount of light spreading in the scintillator.

So, for example, a 600μm CsI layer can have resolution similar to a 300μm thick rare earth screen.

These screens such as gadolinium oxysulfide have the advantage of much lower cost and greater flexibility in that the screen can easily be changed to match the resolution requirements of the application.

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The light generated by the scintillator is absorbed by the photodiodes in the array, creating electrons which are stored on the capacitance of the photodiode itself.

The peak light absorption efficiency of the photodiodes is in the green spectrum, at 550nm wavelength.

Both gadolinium oxysulfide and thallium doped cesium iodide,CsI(Tl), produce their peak light output at this frequency.

The amorphous- silicon photodiodes are typically the “n-i-p” type.

This type of amorphous-silicon photodiode has the advantages of low dark current and a capacitance that is independent of the accumulated signal.

Compared to crystalline silicon photodiodes like those used in CMOS imagers, the dark current in amorphous silicon photodiodes is orders of magnitude less.

So it is not unusual for amorphous silicon flat panel arrays to be capable of more than ten second integration times at room temperature.

The fact that the diode capacitance is independent of signal helps make the detection system linear.

5.4.2 TFT/Photodiode layer

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Figure 5.9:TFT/Photodiode array schematic

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Figure 5.10:A view of a single 127μm pixel.

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Amorphous silicon is not suited to the subsequent signal processing so every column and row of the array is brought to the edge of the glass, where it is connected to a standard crystalline silicon chip by means of a TAB (tape automated bonding) package.

The chips that need to be directly connected to the array, the readout chip and the driver chip, are mounted in these TAB packages.

Figure 5.11:Board-side view of TAB packaged row driver and custom ASIC readout chips.

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The TFT/photodiode matrix is normally scanned progressively, one line at a time from top to bottom.

At the end of each dataline is a charge integrating amplifier which converts the charge packet to a voltage.

At this point the electronics vary by manufacturer, but generally there is a programmable gain stage and an Analog-to-Digital Converter.

One important aspect of the scanning is that the FPD is continuously collecting the entire incident signal; none is lost even during the discharge of the pixel.

The FPD is an integrating detector and the integration time for each pixel is equal to the frame time.

The electronics can be mounted to the side of the array, out of the beam, as is done in higher energy (MeV) applications to protect against radiation damage.

But for diagnostic and interventional procedures, to maintain the best view of the patient, the electronics can also be mounted behind the array and protected by a thin layer of lead.

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Flat Panel Operational Advantages The most obvious advantages of flat panel imagers are size

and weight. The Image Intensifiers Tubes (IIT) are large and bulky. An

FPD with a 12”x16” active area (20” diagonal) takes up less than 25% of the volume of a 12” IIT and less than 15% of that of a 16” IIT.

In addition, the FPD takes the place of not only the IIT, but also the attached image recording devices, including the CCD camera, 35mm Cine camera, and the spot film device.

The result is vastly improved access to the patient in interventional procedures.

In addition to the reduction in size, the weight of the flat panel imager is 60% less than that of the IIT-based imaging chain.

Flat panels also are more economical than an IIT of comparable size considering that we know IIT image quality deteriorates as a simple consequence of everday X-ray use, and thus IIT’s have a relatively short service life.

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Figure 5.12:Signal output comparison vs. X-ray dose between an FPD and an IIT of comparable size.

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FPD Image Quality FPDs have a very direct, short signal conversion

path, with essentially no optics. The result is a very flat, uniform “film-like” image

from edge-to-edge. The ability of flat panel detectors to encompass

multiple X-ray modalities is also a function of their very large dynamic range.

Figure 5.13 shows the signal response of an FPD in the full resolution mode, over the dose range of 1μR to 1.2mR.

Particularly for the amorphous silicon TFT/photodiode technology, the response to entrance dose of the FPD is extremely linear.

The response of the imager deviates from the straight line curve by less than 0.01%.

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Figure 5.13: Signal vs. entrance dose for a 194μm, 40x30 cm FPD.

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Of equal importance is the Signal-to-Noise (SNR) behavior versus dose.

The FPD contributes electronic noise to the image.

However the statistical noise in the X-ray beam is dominant.

The noise in the beam follows Poisson statistics (the noise is equal to the square root of the number of incident X-ray photons)

As shown in Figure 5.14, the SNR of an FPD has a square root dependence on dose, i.e. is X-ray quantum limited over a very large range.

This is an indication that the detector contributes effectively no noise to the image over this dose range.

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Figure 5.14:SNR vs. entrance dose for a 194μm, 40x30cm FPD.

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•The DQE as a function of entrance dose (fluoro and cine range) for an angiographic flat panel imager is shown in Figure 5.15. •This is typical of the indirect TFT/photodiode technology with a CsI scintillator. •Because of the low-loss, high-absorption detection path, the DQE for indirect CsI-based flat panels is the highest available and is more than double that of computed radiography, film screen and CCD based technologies. •Higher DQE translates directly into better imager quality for a given dose. •So with high DQE detection systems, it is possible to get the same image quality as a low DQE system like screen film at a fraction of the dose.

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Figure 5.15:DQE vs. spatial frequency over the fluoroscopic and cine dose range, for a 40x30cm angiographic FPD.

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Examples of Equipments with FPD

Schick technologies intraoral wireless FPD

Model CDR wireless

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Philips Veradius C-arm with FPD

•1.6k x 1.4k image

•Active image area of 28.7 x 26.5 cm

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Philips wireless FPD

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Technology

Medical x ray image sensors