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Thermal imaging is a fast, passive and non-invasive medical imaging modality used to measure and analyze physiological functions and pathology related to the thermal homeostasis and temperature of the body. The technique involves the detection of infrared radiation that can be correlated directly with the temperature distribution of a defined body region. The human body is capable of maintaining a constant temperature that can be different from the ambient temperature. The body temperature is preserved within fine limits typically at 37°C ± 1°C, and it is essential for the normal functioning of the human body.
Change in the body temperature by a few degrees is considered to be a clear indication of probable disease. Thermal imaging has been used successfully in the diagnosis of breast cancer, diabetes neuropathy, and peripheral vascular disorders. It has also been used to detect problems related to gynecology, kidney transplantation, dermatology, heart, neonatal physiology, fever screening and brain imaging. With the advent of modern infrared cameras, data acquisition, and processing techniques, it is now possible to have real-time high-resolution thermographic images.
All objects with a temperature above absolute zero emit electromagnetic radiation, which is known as infrared radiation or thermal radiation. The wavelength of this radiation lies within a range of 0.75–1000 µm. This wide range can be further subdivided into three smaller groups as near infrared – NIR (0.76–1.5µm), medium infrared – MIR(1.5–5.6µm) and far infrared – FIR (5.6–1000µm). According to thermal radiation theory, a black body is considered as a hypothetical object that absorbs all incident radiation and radiates a continuous spectrum. The Stefan–Boltzmann law for total radiation emitted from a perfect black body is given by
R(T) = sAT4
R(T) is the total power radiated into a hemisphere
S is the Stefan–Boltzmann constant (5.67 x 10-8 W m-2 K-4)
A is the effective radiating area of the body
T is the absolute temperature of the radiating surface
For real surfaces,
R(T) = e(T) sAT4
Where e(T) is the emissivity of the emitting surface at a ixed wavelength and absolute temperature T
For a thermal black body emissivity is unity, but for real surfaces emissivity is always less than unity.
The rate of heat loss between two surfaces at T1 andT2 is = A s[ e(T1) T1 4- e (T2) T2 4]
This radiant heat loss forms the basis of thermal imaging.
The emissivity of different human tissues at 40 0C in the infrared wavelength
Black skin (3–12µm)
0.98 ± 0.01
White skin (3–14µm)
0.97 ± 0.02
Burnt skin (3–14µm)
0.97 ± 0.02
Epicardium (fresh:0.5 h) 3µm
Epicardium (fresh:0.5 h) 5µm
Epicardium (9 days at-20 0C)
Infrared emissions from human skin occur between 2 and 20µm with maximum emission around 10µm. The emissivity of human skin is almost constant and its value is 0.98 ± 0.01 for wavelength range of2–14µm. Steketee has found that the emissivity of white skin, black skin, and burnt skin is the same and it is independent of wavelength.
An infrared scanning device is used to convert infrared radiation emitted from the skin surface into electrical impulses that are visualized in color on a monitor. This visual image graphically maps the body temperature and is referred to as a thermogram. The spectrum of colors indicates an increase or decrease in the amount of infrared radiation that is emitted from the body surface. Since there is a high degree of thermal symmetry in the normal body, subtle abnormal temperature asymmetry can be easily identified.
A typical imaging device consists of a system for collecting radiation from a well-defined field of view and a detector that transduces the radiation focused on it into an electrical signal. From 1960to about 1975, images were made with scanning systems that used single IR detectors. The period from 1975 to 1995 saw the commercial development of the SPRITE(signal processing in the element) detector, linear arrays, and 2D arrays. Since 1993 focal-plane array (FPA) detectors have become available.
These are devices of semiconductor-type in which the photon absorption results in the freeing of bound electrons or charge carriers in proportion to the intensity of the incident radiation. In order for the energy gaps to be small enough to allow the detection of radiation beyond 10µm, mixed crystal detectors such as asCdHgTe (CMT) and PbTe have been used.
The parameters by which photon IR detectors are usually specified are responsivity, noise, detective, cut-off wavelength and time constant. Responsivity (R)is the ratio of the output voltage to the radiant input power, expressed in volts per watt. Since the output voltage due to incident infrared radiation is a very small fraction approximately 10-5 of the DC bias voltage across the detector, the responsivity is measured by exposing the detector to chopped radiation from a calibrated source at 500 K and measuring the alternating voltage component at the chopping frequency. Responsivities (R,500 K) are typically 104-105 V W-1 for CMTdetectors operated at -196°C. Noise, together with responsivity determines the ability of a detector to detect small input signals. This is specified in volts per hertz at one or a number of frequencies or as a noise spectrum.
Detectivity(D) is given by:D = 1/NEPwhere NEP is the noise equivalent power, which is the RMS value of the modulated sinusoidal radiant power that falls upon a detector that will give rise to an arms signal voltage (Vs)equal to the rms noise voltage (VN) from the detector. For many photon detectors, the NEP is directly proportional to the square root of the area of the detector and it becomes appropriate to use a normalised detectivity D* given by* = DAd1/2 (if) = (Vs/ Vn)[Ad(if) ]1/2/ WwhereAd is the area of detector, if is the frequency bandwidth of the measuring system,W is the radiation power incident on the detector (RMS).
D* varies with the wavelength of the radiation and the frequency at which the noise is measured. As a figure of merit, D*enables a theoretical maximum detectivity to be calculated that would apply when performance is limited only by noise due to the fluctuation of background radiation. The detector time constant G is the time between incident radiation being cut off and the output of the detector dropping by 63%. Typical time constants range from a fraction of a microsecond for CMT detectors to a few microseconds for InSb detectors.
The detector forms a part of an imaging system whose performance is generally specified in terms of temperature resolution, angular resolution, and field of view. Temperature resolution is a measure of the smallest temperature difference in the scene that the imager can resolve. It depends on the efficiency of the optical system, the responsivity, and noise of the detector and the SNR of the signal-processing circuitry. Temperature resolution can be expressed in two ways: noise equivalent temperature difference (NETD), which is the temperature difference for which the SNR at the input to the display is unity and minimum resolvable temperature difference (MRTD), which is the smallest temperature difference that is discernible on the display. Most of the medical thermography has been carried out with systems that have MRTD between 0.1 and 0.3 K. Angular resolution is typically 1–3 mrad but can be as small as 0.5 rad.
In an imaging system that is capable of transmitting and focusing infrared radiation, the scene is viewed by an optical lens. A high value of IR refractive index is advantageous in the lens design but materials that have high refractive indices tend to have low transmittance. This high reflective loss may be eliminated by antireflection coatings, which increase the transmittance up to 95% – 97% for a given wavelength range. Germanium and Silicon are used frequently for IR optical components.
Advances in IR technology have resulted in improvements in the resolution of imaging systems and designers of imaging systems struggle to develop matching optical systems. This has led to a search for achromatic IR-optical materials and the use of chalcogenide glasses. This selenium based glasses combine well with other IR optical materials and provide high-resolution optical material operating in the 3–5 and 8–12 µm atmospheric windows.
First and second generation scanning systems used configurations of lenses, rotating prisms, rockingmirrors or rotating multisided mirror drums. The precise design was dependent on commercial features and the functional purpose of the imager. Single-element detectors have the advantage of simplicity, both electronically and mechanically. There are two advantages to replacing a single detector with a multi-element array of n similar detectors. There is a decrease in noise level since the signal increases in proportion to n whereas noise increases in proportion to n1/2 and the higher scan speeds that can be obtained with array detectors make this instruments important for investigating rapid temperature changes. In a clinical context, high-resolution real-time imaging allows a precise focusing on the skin surface and continuous observation of thermal changes so that transient or dynamic studies can be made on patients.
A significant development in the imaging technology was the SPRITE detector. The SPRITE CMT detector performs the delay and adds functions within the element so that a single detector replaces a linear array of detectors. In a conventional in-line array, the signal from each IR-sensitive element is pre-amplified and then added to the signal that is generated in the adjacent element. In the SPRITE, the individual elements are replaced by a single IR strip mounted on a sapphire substrate. It requires only one amplifier channel and has the optimum gain at high speeds. An eight-elementSPRITE detector is equivalent in performance to an array of at least 64discrete elements but requires far fewer connections. By arranging the detectors in a stack, outputs can be stored in parallel in-line registers serially combined to a TV compatible display rate.
Since1993 much of the earlier technology that employed single detectors, linear arrays and SPRITE detector arrays have been replaced by the development of starting-array detectors. These third generation cameras offer higher temperature resolution images in real time. The absence of a scanning mechanism means that all FPA solid-state cameras are very compact and quiet in use. The scene is viewed through a lens and when appropriate, a filter can be incorporated to view IR above a specific wavelength. FPA arrays have been constructed from a number of different materials including InSb, PtSi, CdHgTeand InGaAs. Complex scanning systems are no longer required and the inherent simplicity of the staring-array detector together with advances in micro-cooler technology has resulted in the manufacture of very compact, high-performance string-array systems.
Ina scanning system, the detector or each pixel of the detector only sees the object for a very short time and this reduces the amount of energy collected. In FPA systems, the scanned detector is replaced by an array of detector cells, one for each pixel staring constantly at the object being imaged. To increase the number of detectors inside the sensor vacuum Dewar, most quantum detector arrays operate in photovoltaic mode. The detectors can be fabricated on a substrate as p–n junctions using integrated circuit techniques with very high packing density. There are two ways of constructing FPA detectors.
Monolithicdetectors are easier and cheaper to construct because both the IR-sensitive material and the signal transmission paths are on the same layer. On the contrary, in the case of hybrid FPAs, the detector is on one layer and the signal and processing circuitry is on another layer. The advent of micron and sub-micron silicon technology has led to the manufacture of complex signal conditioning electronics and multiplexers integrated onto a silicon chip. This, in turn, is incorporated directly behind the IR detector within the vacuum encapsulation. The problem of non-uniformity of detector response in FPAs is addressed by using digital signaling electronics and computing technology to match all channels.
UncooledFPA detectors based on the principle of the bolometer have also been developed. These devices consist of a sensitive area whose electrical resistance is strongly dependent on its temperature. The absorption of the incident thermal radiation changes the temperature of the sensitive area and the change in the measured electrical resistance results in a signal proportional to that radiation. The disadvantage of this type of thermal detector is that they react relatively slowly compared to the response of photon detectors. However, such detectors respond fast enough to work well in FPA systems where response requirements are in the millisecond range.
Staring-array technology has also been applied to quantum-well-type IR photodetectors(QWIPs). These devices are built to have a quantum well with only two energy states, the ground state, and the first excited state. The excited state is arranged to be near the top of the well so that it can detect light photons. By alternating layers of the wall material such as GaAs and the potential barrier, it is possible to control the characteristics of the QWIP so that it will respond to a particular wavelength of radiation. Normally, QWIP detectors are designed to detect radiation in the 8–9 µm range.
FPA-based systems requiring detector cooling have also been developed. In these cameras, the matrix of detector cells is fashioned from InSb or from PtSi and must be cooled to 80 K for optimal use as a thermal imaging device. PtSi has reliable long-term stability but it has low quantum efficiency. It is sensitive to radiation in the range of 1–5 µm. Cooling of FPA detectors is usually accomplished by a Stirling cooler. Typically, matrices used in clinical imaging either 320 x 240 pixels or 640 x 320 pixels but for research purposes, detectors with arrays of 512 x 512 pixels and 1024 x 1024 pixels have been developed.
The image quality of FPA detectors that are used clinically is superior to that of the previous scanning systems. Image capture and image processing are easier and faster. Clearly, the use of an uncooled device for dynamic imaging of patients in a ward or clinic is advantageous.
The pyroelectric effect is exhibited by certain ferromagnetic crystals such as for barium titanate and triglycine sulfate (TGS). When exposed to a change in radiance, these materials behave like capacitors on which electrical charge appears. The magnitude of the effect depends on the rate of temperature change in the detector, so the sensor does not respond to a steady flux of radiation. Pyroelectric detection has been developed as a cheaper alternative to photon-detector based systems. Although pyroelectric detectors were originally incorporated into systems employing mechanical scanning devices to construct an athermal image, most values for clinical work came from the development of pyroelectric vidicon camera tubes.
In pyroelectric systems, the scene is panned or modulated by a rotating disc and the IR radiation enters the vidicon tube by means of a germanium IR transmitting lens (8–14 µm) which focuses the image of the thermal scene onto a thin disc of TGS pyroelectric material. At the front of the TGS target, there is an electrically conducting layer of material that is chosen to be a good absorber of thermal radiation. The target is scanned ina TV raster by the electron beam of the vidicon tube and the image is displayed on a TV monitor. The latest pyroelectric imagers are based upon FPAs that use the pyroelectric effect in ceramic barium-strontium titanate. Multi-pixel arrays having an NETD of 0.5°C have been developed largely for industrial use and surveillance purposes.
Rheumatoid arthritis is a chronic inflammatory disorder that affects the joints. It results in over perfusion of the tissue and a consequential increase in the skin temperature. Thermal imaging provides objective, quantifiable and reproducible measures of the intensity and the extent of joint involvement. It can be distinguished between deep-seated inflammation and more cutaneous involvement. By standardising the conditions and cooling the peripheral joints so that the skin is within a specific temperature range (26°C–32°C for the lower limbs and 28°C–34°C for the joints of upper limbs), it is possible to quantify the thermal pattern in the form of a thermographic index on a scale from 1.0 to 6.0, in which healthy subjects are generally found to be less than 2.5 while inflammatory joints rise to 6.0. This quantitative analysis is a very effective means of assessing the efficacy of anti-inflammatory drugs used in the treatment of rheumatic conditions.
Raynaud’s disease is a medical condition in which the spasm of the fingers’ arteries causes episodes of reduced blood flow, provoked by cold or emotional stress. Through thermal imaging, the severity of the disease can be quantified and consecutive attacks can be compared with each other. Due to the difference in the underlying disease processes, the primary and secondary forms of Raynaud’s disease can be differentiated by thermal imaging, which provides valuable information for further diagnostic procedures and individualized management of the disease and also has prognostic value.
Knee osteoarthritis is the degenerative disease of the tissues of the knee joint, accompanied by an inflammatory process of varying degrees. Morphological changes observed by imaging methods can be detected only after a long period of time from the onset of signs and symptoms of the disease, even when sensitive imaging methods are used. For these reasons, the evaluation of therapeutic efficiency and the decision on the continuation of a certain therapy or its replacement by an alternative therapy cannot be based on the results of the traditional imaging modalities during the early modifiable course of the disease.
The patellaphysiologically represents a cool spot with a characteristic shape in the thermal imaging, because its thick bone plate prevents the dissipation of the heat produced by the knee joint through the patella and thus the heat is dissipated around the patellar margin, which can be detected by the presence of a slightly warmer band that surrounds the patella. In case of inflammatory processes of the knee, the normally cool spot that represents the patella with the surrounding slightly warmer band becomes distorted or disappears completely and the temperature of the skin covering the inflamed knee tissues rises. Even in advanced osteoarthritis that is detectable in X-rays, the increased temperature of the skin covering the patella correlates with the severity of the radiographic changes. Quantitative assessment of pain-related thermal dysfunction by thermal imaging is utilized in other parts of the body as well.
In plastic and reconstructive surgery thermal imaging is used in many ways. It is an excellent diagnostic tool to identify the dominant perforator vessels before free flap surgery, which helps in pre-operative planning. It is an exceptional method to monitor the perfusion of the free flap after connecting its vessels-arteries and veins to the site of reconstruction intraoperatively. In the post-operative period itis a sensitive, valuable method to assess the free flap in difficulty and to decide if the clinical symptoms are related to problems with flapper fusion or are due to other causes such as infection.
Minute changes of less than 0.01 K in the temperature of the cerebral cortical surface can be detected by thermal imaging. It has led to the proof of concept study of measuring the cortical cerebral perfusion by the cold saline technique. A small amount -10mL of ice-cold saline was administered as a bolus into a central vein and subsequent changes in cerebral cortical temperature have been recorded by infrared video thermography and analyzed by the principal component analysis (PCA) in patients who were operated on for their cerebral pathologies such as ischemic stroke, brain tumor etc. It has been shown, that the method is able to differentiate between cortical regions with good or poor perfusion.
Infrared imaging and image analysis have been introduced as a powerful tool for deception detection around 2000 by the demonstration of its ability to detect facial stress patterns at a distance. The breathing cycle can be monitored by thermal imaging based on the difference in temperature between the exhaled air and the ambient temperature. The cardiac pulse wave can also be monitored by infrared thermography in locations, where large arteries travel close to the skin surface. Replacement of the traditional polygraph testing by non-contact infrared video thermograph recordings and their multifaceted analysis makes it possible to test a large number of individuals potential signs of their deceptive behavior.
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