Imaging methods employ different forms of energy to probe and detect anatomy and biological processes. Figure 14–2 presents these energy forms together with an overlay of exogenous imaging agents. In the following sections, the dominant imaging modalities used in cancer are reviewed in terms of the process of signal formation, as well as a synopsis of their more interesting applications in oncology.
Images can be developed from endogenous signals or from exogenous signals induced by the introduction of contrast agents or molecular probes. The detection of these signals is through either electromagnetic or acoustic energy transfer. Imaging agents are designed to either produce the detected signal (eg, radiolabeled PET/single-photon emission computed tomography [SPECT]) or to alter the interaction between the object and the applied energy (eg, high atomic number materials such as iodine, barium, or xenon in radiography and CT, paramagnetic agents in MR, bubbles in ultrasound). Optical imaging operates at energies corresponding to biological processes and can therefore provide insight into active biological processes (eg, detection of bioluminescent signals). Similarly, MR offers insight into the chemical activity in the body by exploiting effects related to the impact of chemical milieu on nuclear magnetic resonance to create image-like maps of these effects. This is referred to as MR spectroscopy imaging (MRSI). Chemical milieu can also be probed through the exchange of protons between specific chemical species (CEST).
14.3.1 X-ray–Based Systems: Radiography and Computed Tomography
X-ray imaging is based upon the differential attenuation of x-rays within different tissues in the body. The discovery of x-rays by Roentgen in 1895 revealed the value of noninvasive imaging and its clinical applications were immediate. Despite 100 years in advancing this technology, the main methods of x-ray generation and detection have not changed substantially: a vacuum x-ray tube and a 2-dimensional (2D) detector on either side of the subject have remained the central elements. Although the chest x-ray is still a useful tool, x-ray–based CT has become a standard technology to detect and stage cancer since its invention in the 1960s by Hounsfield and Cormack (for which they shared the 1979 Nobel Prize in Physiology or Medicine). The basic process for image generation is shown in Figure 14–3A. Briefly, a highly efficient 2D array of x-ray detectors is located opposite a powerful x-ray tube operating in excess of one hundred thousand volts (~120 kVp). The gantry rotates through 360 degrees over a period of less than 0.5 seconds while acquiring hundreds of digital radiographs of the patient. A computerized process of digital filtering and back-projection allows an estimate of the x-ray attenuation coefficient of each voxel in a "slice" (1 to 5 mm thick) through the patient to be estimated (Kak et al, 1988). Modern CT scanners can acquire multiple slices simultaneously and images of the entire body can be acquired by moving the patient through the CT scanner during rotation. Computerized postprocessing of the individual voxels then allows reconstruction of an image in any plane desired.
Imaging technologies used in oncology. A) In a modern CT imaging system, a multirow detector and high-power x-ray source rotate about the patient at high speed (>2 revolutions per second). Thousands of "projections" are collected as the patient advances through the scanner and computers are used to "reconstruct" multiple imaged slices representing the attenuation characteristics of the patient. This is presented in grayscale form as the CT image. B) An MR image signal has many forms, but relies largely on the excitation and relaxation of a population of field-aligned magnetic dipoles associated with protons (hydrogen nuclei) in water. MR imaging systems consist of a large superconducting magnet that maintains the static field (typically 1.5 Tesla), that causes the field alignment of the magnetic dipoles of hydrogen nuclei of the water molecules, and a set of gradient-inducing coils that manipulate the magnetic field throughout the volume. The radiofrequency transmit and receive coils are responsible for perturbing the hydrogen nuclei and then recording their relaxation back to the ground state in the presence of the magnetic field. C) PET imaging is often used in conjunction with a CT system (called a PET-CT scanner). PET image formation is achieved through detection of positron-emitting decay events that ultimately produce pairs of 511 keV photons at approximately 180 degrees from each other. This image is then superimposed on the CT image generated at the same time by the dual-purpose scanner. D) Ultrasound (US) imaging systems exploit variations in acoustic impedance within the body to generate images. Ultrasound waves are reflected at boundaries between tissues of differential impedance (eg, fat, muscle, bladder wall). In this figure, an axial ultrasound image of the prostate is shown as generated by a transrectal ultrasound probe (see illustration).
The CT image signal is measured in Hounsfield units (HU), a linear measure of the attenuation coefficient of the tissue relative to water (water is 0 HU and air is–1000 HU). Images formed at typical imaging doses would have approximately 1% noise (10 HU) and percent differences in HU between tissues such as fat and muscle are only approximately 5% (corresponding to a differential of ~50 HU). It was originally hoped that quantifying the CT signal could be used to classify normal tissue and disease, but the lack of specificity of the CT signal has resulted in little progress on this front. However, the recent development of dual-energy CT imaging systems provides a much higher fidelity in characterizing tissues, and these systems will revitalize the concept of tissue classification (Gupta et al, 2010). High-atomic-number contrast agents (eg, iodinated molecules) are used routinely in CT imaging to increase the contrast-to-noise ratio of various structures and are used in bolus studies with fast (2 to 3 images per second) repetitive scanning to study the permeability of tissues perfusion (Miles, 1991). Xenon gas has also been explored as an agent that is inhaled to assess lung ventilation and as an agent to measure tissue blood flow in hepatocellular carcinoma (Murakami et al, 2004).
CT has the advantage that it achieves high spatial resolution (<1 mm), soft-tissue discrimination, is highly reproducible, and can be employed quantitatively for measuring tumor size and detecting response. However, the image quality in terms of SNR is related to the applied dose of ionizing radiation—typical CT imaging doses range from 1 to 10 cGy (see Chap. 15, Sec. 15.2.2 for definition of radiation dose). As the first soft-tissue volumetric imaging modality, CT transformed oncological imaging and is used routinely for detection and diagnosis of cancer, as well as for monitoring patients that have undergone therapy. The clinical impact is seen in the management of the patient through noninvasive cancer staging (see Chap. 10, Sec. 10.3.1 for brief description of TNM staging), which has led to a reduction in the frequency of exploratory laparoscopic surgery. Characterization of regional extent of disease is undertaken using pretreatment, posttreatment, and follow-up CT imaging. CT also has a major role in directing therapy, including its use in directing radiofrequency ablation of liver metastases by interventional radiologists, assisting surgeons in head and neck surgery, and delineating lung tumors for image-guided stereotactic radiotherapy. There is widespread use of CT-guidance in interventional radiology and radiation therapy.
CT is used to screen for lung cancer in higher risk populations (see Chap. 22, Sec. 22.3.3) and randomized clinical trials have demonstrated its ability to improve outcome for people with lung cancer (Aberle et al, 2011). The value of CT screening for cancer in other sites has not been established and there are concerns about potential health effects and secondary malignancies relating to the use of CT scanning for screening the general public (Brenner et al, 2007; see Chap. 16, Sec. 16.8).
Recent interest in characterizing the vascularity of tumors has spurred the use of dynamic CT imaging (multiple slices/volumes per second) in which CT scanning is undertaken with a bolus intravenous (IV) injection of a low-molecular-weight-iodinated agent to characterize the vascularity and permeability of cancer lesions as it passes through the patient's tissues. This is referred to as perfusion CT, functional CT, or, more recently, dynamic contrast-enhanced computed tomography (DCE-CT). One application of these methods is to detect response of hepatocellular carcinoma to antiangiogenic therapies. DCE-CT is sensitive to vascular changes that occur before conventional size-related measures of tumor response (Jiang et al, 2011).
14.3.2 Magnetic Resonance Imaging
The MR image signal arises from the following 5 concepts (see Fig. 14–3B): (a) a large ensemble of nuclear magnetic moments or "tiny bar magnets" (eg, hydrogen nuclei in water) are contained within the human body; (b) an applied, static magnetic field (eg, 1.5 Tesla [T]) sets a slight bias in their orientation—the higher the field, the larger the bias; (c) the application of an external radiofrequency field to perturb the moments from this bias; (d) the surrounding chemical and physical microenvironment impacts the time required (T1) to return to alignment (relax) with the static field; and (e) the ability of the ensemble to induce a measureable current in a detecting antenna (conducting loop) exterior to the object. Protons (and other atomic nuclei possessing a magnetic moment) within tissues oscillate, or precess, in this magnetic field (B) at a frequency (w) given by w = γB where the proportionality constant γ is called the gyromagnetic ratio; γ is specific for each nucleus and depends on the magnetic moment of the nucleus.
Damadian (1971) first proposed that the rate of relaxation of protons in water could distinguish normal from tumor tissue and demonstrated that tumors (sarcomas and hepatomas) had differing T1 relaxation times that were 1.5 to 2 times longer than in normal tissues in the same animal. These differences in relaxation times are a reflection of the different environments in which the protons relax to alignment with the static field—the presence of cancer alters this environment compared to that found in normal tissue.
The development of methods to generate images of nominal relaxation times followed from the work of Lauterbur (1973) and Mansfield (Mansfield and Maudsley, 1977) for which they received the Nobel Prize in Medicine in 2003. MR imaging (MRI) systems are now widely used in cancer detection and diagnosis. Figure 14–3B demonstrates the central components: A static magnetic field (typically 1.5 Tesla or 15,000 times the earth's field) is generated by a superconducting magnet; the patient is placed within the magnet; gradient coils are used to create slight differences in magnetic field across the patient to encode for location; and, finally, a pair of antennae is responsible for exciting the nuclei and detecting the electromagnetic signal that they induce as they return to their ground states. Spatial information that allows the formation of images is obtained by slightly varying the applied magnetic field across the body in 3 orthogonal directions using the gradient coils.
In a 1.5-Tesla magnetic field, the precession frequency of water protons is 64 MHz. Radiofrequency (RF) pulses of energy applied at this frequency alter the angle at which the protons are precessing around the magnetic field lines. Once a 90-degree pulse is switched off, the protons gradually return to alignment with the static field. The time taken for the return to alignment (the T1 signal, also called the spin-lattice relaxation time) is one metric of the local chemical and physiological environment of the protons and can provide contrast between tissues. Similarly, the phase of their precession can be synchronized within the transverse plane (through a 180-degree pulse). When the RF is switched off, dephasing gradually occurs and the time to reduce the transverse magnetization is referred to as T2. This dephasing is associated with both molecular effects and inhomogeneity in the magnetic field; T2 is used commonly to distinguish soft tissues and is called the spin-spin relaxation time.
Because complete isolation of T1 and T2 signals is challenging, images are typically T1 or T2 "weighted" depending on the RF pulse sequence applied. Specifically, by varying the repetition time (TR) between RF pulse cycles and the time to sample the resulting signal or the echo time (TE), it is possible to "weight" the image. There are 3 basic weightings used in clinical practice: T1, T2, and proton-density weighting. In general, T1 weighting provides anatomical detail, while T2-weighted images give elevated signal for tissues with a higher content of free unbound water and are essential in imaging inflammation or "neoplastic" tissue. Proton density simply reflects the density of water protons available for signal production in the voxel.
MR can be used to explore other tissue parameters, such as changes in local transport of water in tissue that may reflect cellular sensitivity to treatment. This can be characterized by exposing excited nuclei in tissue to variations in a magnetic field and examining the rate of signal loss; this is referred to as apparent diffusion coefficient (ADC) imaging (Le Bihan et al, 1986). Figure 14–4A illustrates the process used in MR to estimate the diffusion of water within a voxel (Hagmann et al, 2006). In brief, a conventional spin-echo sequence is modified by applying 2 gradient pulses before and after the 180-degree pulse to encode the degree of proton mobility (ie, diffusion) into a loss in recovered signal as a result of transport-induced imperfections in rephasing of the spins. Figure 14–4B presents diffusion-weighted images of a patient before and after treatment for lymphoma.
A) The ADC (or diffusion-weighted imaging [DWI]) imaging technique seeks to measure the diffusion of water within a voxel by exposing the excited water protons to a spatial gradient during their dephasing to encode for their diffusive transport. This exposure is done before and after a 180-degree spin-echo pulse. Voxels that contain spins exposed to differential fields because of diffusive transport will have a reduced echo following the 180-degree rephasing pulse. Those that are stationary will be refocused to within normal T2 losses. B) DWI is also of interest in assessing total cancer burden, as these techniques can also be used to image the entire body. In their review of DWI in oncology, Padhani et al (2011) demonstrate DWI in the assessment of pretreatment disease burden and its response to chemotherapy in a patient with Hodgkin lymphoma. The 2 panels show pre- and posttreatment DWI images. (Reproduced with permission from Padhani et al, 2011.)
Exogenous agents can be applied to alter the MR signal. Spin-lattice relaxation time (T1) and spin-spin relaxation time (T2) may be shortened considerably in the presence of paramagnetic species (eg, gadolinium), which have unpaired electrons. Stable agents that contain gadolinium are used in the clinical setting (eg, gadopentate dimeglumine or gadolinium-diethylenetriamine pentaacetic acid [Gd-DPTA]; gadoteridol [Gd-HP-D03A]), and there is growing interest in iron oxides, for detection and characterization of nodal disease (Harisinghani et al, 2006) or for use in cell "tracking" studies in animal research (Heyn et al, 2006). New MR contrast agents are being developed, including some that rely on the rapid exchange of protons between molecular environments to describe the chemical milieu (Sherry and Woods, 2008) and some that can increase the polarization of specific nuclei to enhance the MR signal by several orders of magnitude (Golman, Ardenkjaer-Larsen et al, 2003).
Magnetic resonance spectroscopy imaging (MRSI) involves the extension of nuclear magnetic resonance (NMR) techniques employed in chemistry to the concepts of imaging. Spectra associated with the chemical shift in resonance peaks for various biomolecules in small regions of interest can be acquired on MRI systems. Collecting a number of spectra in adjacent regions in a rectilinear array (Fig. 14–5A) and generating a coarse image is referred to as MRSI or chemical shift imaging (CSI). MRSI can be applied to protons in water or other nuclei with a magnetic moment such as 31P, 13C, and 19F. Figure 14–5B illustrates the nature of the spectra produced on a clinical 1.5-Tesla MR scanner when imaging the prostate; the poorly resolved spectra are reduced to ratios of peaks to form a color-coded image of disease burden. Adoption of higher field (3 Tesla) MR scanners will enable greater spectral separation of the peaks and may lead to accelerated adoption of these techniques (Glunde et al, 2011).
Magnetic resonance spectroscopy (MRS) offers the potential for metabolic imaging by detecting molecular environment-induced frequency shifts in the resonance of nuclei. A) A set of MR spectra acquired in a rectilinear array and overlaid on the conventional MR image of the prostate (1.5 Tesla with endorectal coil). The spectra contain a citrate peak present in normal and disease tissues, while the elevated choline peak corresponds to disease (B). C) Spectral analysis in these studies consists of calculating ratios of signal over specific frequency intervals or peak heights estimated from peak-fitting algorithms (eg, Cho/Cit ratios). These ratios are then converted to a color-coded pattern and overlaid on the MR image to identify regions of elevated disease burden.
The excellent soft-tissue discrimination of MR makes it well-suited to oncology and MRI has a role in the diagnosis, staging, and management of many solid cancer. For example, MR has become the dominant imaging method for cancers of the central nervous system where the high T1 contrast and sensitivity of T2 to changes in edema delineate the extent of disease and allow understanding of the patterns of spread. Its role in evaluating metastatic lesions in the spinal column, including those requiring urgent treatment because of spinal cord compression is definitive. MRSI is also employed in neurooncology as an additional classifier of disease prior to surgical intervention (Hollingworth et al, 2006). Although use of MR in breast screening or directing surgery is controversial, it is useful as a screening method for women at high risk (~20% probability) of developing breast cancer (Lehman et al, 2009). Dynamic contrast-enhanced magnetic resonance (DCE-MR) (conceptually equivalent to DCE-CT) is emerging as a biomarker of drug activity in breast cancer (Moon et al, 2009) and is being used in clinical trials that assess response of disease to antivascular agents (Yankeelov and Gore, 2009) and to radiation therapy (Cao, 2011). There is a growing interest in the development of dedicated MRI systems to guide cancer therapy; this includes the direction of high-intensity focused ultrasound (HIFU), neurosurgery, and MR-guided radiotherapy (Lagendijk et al, 2008). Until recently, the relatively slow acquisition times have limited the use of MRI in sites influenced by motion (eg, chest), but this is changing rapidly with the development of faster imaging techniques that employ multiple channels. The progression to higher field strength (3 Tesla) offers increases in SNR and spatial resolution, and MR is attractive in that there is no exposure of subjects to ionizing radiation and hence to risk of second (radiation-induced) malignancies (see Chap. 16, Sec. 16.8).
14.3.3 Single-Photon and Positron Emission Tomography
Unstable nuclei that emit high-energy gamma-rays or positrons provide a powerful tool in probing the nature of cancer and contribute to its management. Molecules containing such nuclei can be injected into the body and accumulate through a variety of processes that reflect different metabolic aspects of the disease state. Single-photon emission computed tomography (SPECT) imaging operates on the principle of emission of gamma-rays in the range of 100 keV that are detected by a collimated crystal (gamma camera) that encodes for direction and location in the field of view. The gamma camera images are then used to reconstruct an estimate of the distribution of the gamma-ray emitters in the body. SPECT is heavily utilized in cancer for the staging and assessment of cancer progression in the form of a "bone scan." In this technique, the patient is injected with a small amount of radioactive material such as technetium-99m (99mTc)-labeled medronic acid (a bisphosphonate) and then scanned with a gamma or 3-dimensional (3D) SPECT camera. The accumulation of the medronic acid in regions of bone remodeling is a sensitive detector of metastasis to bone.
In contrast to SPECT, PET employs radioisotopes that emit a positron upon decay. For example, a radioactive isotope of fluorine (18F) emits a positron that annihilates through interaction with an electron to produce a pair of annihilation photons (511 keV each), which are emitted at approximately 180 degrees to each other. Figure 14–3C illustrates the coincident detection of the event in detectors distributed on multiple rings around the patient. Large numbers of these coincident events are then used to generate (or reconstruct) an image of the spatial distribution of the annihilation events, which represents the distribution of the 18F-labelled agent. The most widely used PET agent is a radiolabeled sugar that becomes trapped within cells that have active glucose metabolism, as is the case for malignant tumors (see Chap. 12, Sec. 12.3.1). 18F-Fluorodeoxyglucose (FDG) is injected intravenously into the body and allowed to circulate and metabolize for approximately 45 minutes. The patient is then positioned in the PET scanner and images are collected to image the whole body for regions of elevated uptake.
There are advantages and disadvantages to the SPECT and PET approaches. The ability to integrate radioisotopes directly into a molecule of interest minimizes impact on the pharmacokinetics of the agent itself. Moreover agents can be designed to target many aspects of tumor cell metabolism and the tumor microenvironment. However, the half-life of the probe needs to be selected for the specific objective: the radioactive half-life needs to be longer than the half-life of the physical process that one is interested in characterizing, but it shouldn't be much longer, or the radiation dose delivered would be excessive. Some positron-emitting radioisotopes (15O, 11C) can be readily integrated into various molecules but have very short half-lives and require production in a cyclotron that is adjacent to the radiochemistry laboratory and the imaging suite. 18F, the most commonly used positron-emitting radioisotope, has a half-life of 110 minutes and therefore can be shipped from remote sources within a few hours with sufficient activity for imaging purposes. 99mTc is the most commonly used SPECT isotope and is a daughter product of a molybdenum generator that can be located within the local SPECT laboratory.
The development of combined PET-CT and SPECT-CT systems was motivated by the advantages of using CT to correct for attenuation by the patient's anatomy in reconstruction of the image (Townsend et al, 2003), but it is now recognized that the additional imaging information provided from the CT is also beneficial. FDG-PET is employed in detecting regional disease and metastatic lesions in people with cancer; it increases the accuracy of staging and thereby promises to lead to better outcomes (Al-Ibraheem et al, 2009). The role of FDG-PET images for target delineation in radiotherapy is a topic of intense research with its value varying with disease site (Gregoire et al, 2007). In addition to imaging metabolism, PET has been employed to characterize the degree of hypoxia in solid cancers (see Chap. 12, Sec. 12.2.1 and Chap. 16, Sec. 16.4). F-misonidazole (FMISO)-PET and 18F-fluoroazomycin arabinoside (FAZA)-PET are hypoxia localizing agents that become trapped in cells as a result of reduction in the absence of oxygen, so that accumulation occurs only in cells that have low oxygen concentration (Padhani et al, 2007; Krohn et al, 2008). Figure 14–6A illustrates the process of trapping of FMISO in the cell under conditions of hypoxia. The dynamics of agent delivery can also be used to examine perfusion of the tissues (Fig. 14–6B; Thorwarth et al, 2005). Numerous cancer-related PET agents are in development, but FDG remains the only one in routine use: Challenges to bringing additional agents to the clinical domain include cost, the requirement for validation studies, regulatory constraints, and difficulty integrating the new image-based information into clinical use.
A) Hypoxia results in the accumulation of FMISO in the cell through a process that is dependent on reduction of the nitro (NO2) group by 1 e– nitroreductases. If O2 is not present, the tracer is sequentially further reduced to an alkylating agent and is bound in the cell. In the presence of oxygen, the initial reduction step is back-oxidized to recreate the original molecule, which can diffuse out of the cell again (Mason et al, 2010). B) This illustrative figure adapted from Padhani et al (2007), with permission, shows an FDG-PET image (bottom left) with increased uptake in both the oropharyngeal tumor (arrow) and in the left neck nodal metastasis (asterisk). The FMISO-PET images (bottom right series) were selected from the dynamic acquisition after 1 minute, 30 minutes, and 240 minutes. The early distribution (1 minute) shows hyperperfusion in the region of the primary tumor and metastasis. However, only the left neck nodal metastasis is shown to retain FMISO (and hence is suspected of containing hypoxic cells) after 240 minutes.
14.3.4 Ultrasound Imaging
Ultrasound imaging utilizes the variations in acoustic impedance in the different tissues of the body to generate anatomic images. Figure 14–3D illustrates a modern ultrasound probe configuration for transrectal ultrasound of the prostate. Piezoelectric crystals are capable of generating very high-frequency acoustic waves (ultrasonic) in the range of 1 to 20 MHz and correspond to short wavelengths given the speed of sound in tissue (~1500 m/s). The ultrasound image is formed through careful manipulation of the acoustic source and detection of the reflected acoustic pressure wave. Although ultrasound is limited in its depth of penetration (1 to 15 cm), it has many useful advantages including low cost, adaptability to small probes for directing minimally invasive biopsy procedures (eg, endobronchial ultrasound, prostate biopsy) and use in tissues that are relatively homogenous (eg, liver imaging). In addition to structural information, ultrasound also offers accurate assessment of flow rates in large vessels using the Doppler phenomenon, in which, the relative motion of the blood induces frequency shifts in the reflected sound waves. Ultrasound can also be applied with contrast agents, such as, microbubbles (Wilson and Burns, 2001) that reflect the sound waves producing very-high-contrast signals. This technique has been used to study tissue perfusion in organs and neoplastic lesions (Delorme et al, 2006). Researchers are also applying molecular imaging approaches to specifically target the bubbles to endothelial cell surface receptors to allowing imaging of the density of such receptors in normal tissue or tumor blood vessels (Caskey et al, 2011).
In the clinic, ultrasound is used widely in the detection, diagnosis, clinical staging, and treatment of cancer. Ultrasound has proven useful in the characterization of lesions in the breast, liver, and kidney, and recent use of ultrasonic contrast agents such as bubbles has improved this performance substantially (Quaia et al, 2006). In addition to diagnosis, ultrasound is used for directing biopsies, providing information in the surgical management of prostate or pancreatic lesions, and is also employed for localizing structures during radiation therapy. For example, transabdominal ultrasound imaging is used in daily radiation treatments to localize the prostate and adjust the patient position just prior to treatment. In the context of brachytherapy of prostate cancer using permanent radioactive seed implants, transrectal ultrasound is used to localize both the gland and the seeds as they are placed in the gland by the oncologist.
Optical imaging continues to play an expanding role in characterization and management of cancer. The development of endoscopic systems has been accelerated by advances in fiberoptic and camera technologies that permit high-quality optical imaging to be applied within the body's orifices and luminal structures (eg, colonoscopy, bronchoscopy, and cystoscopy). High-definition video capture, integration of channels for biopsy/resection, and the development of navigation technologies is increasing the application of these systems. These systems can collect light over a wide spectrum from near infrared to blue or operate on a narrow band. In addition, they can be used with different light sources ranging from white light to higher-frequency stimulating sources to generate fluorescence. The endogenous absorption of light in tissue provides significant signal for the detection and characterization of disease by these methods. Although optical imaging suffers from scattering and absorption within the body, it has several advantages, including the high detection efficiency of optical detectors, the abundance of light photons that can be generated and interact without harm to tissue, and the potential to support molecular imaging. The potential for optical imaging to increase the sensitivity of detecting lesions on luminal structures (eg, colon, glottis, esophagus) has resulted in substantial research in this area. The detection of autofluorescence demonstrating increased sensitivity to cancer lesions in the oral cavity is a very promising example (Poh et al, 2006). This is illustrated in Figure 14–7 where (a) white light, (b) autofluorescence, and (c) narrowband imaging each provide increasing detail in the characterization of a gastric adenocarcinoma.
Endoscopic techniques continue to evolve as demonstrated in this series of images from Filip et al (2011). Endoscopy of an early gastric adenocarcinoma at the level of the gastric angle reveals (A) an irregular ulcer visualized in white-light endoscopy and (B) autofluorescence images showing in magenta the neoplastic margins and a larger lesion extension, as compared with white-light endoscopy. Further advances in resolution and selection of a narrow wavelength band (C) shows a modified pit pattern, with irregular and distorted vascular pattern in the center suggesting high-grade dysplasia/early cancer.
Endoscopic imaging is a heavily used imaging modality in cancer detection and in the context of colon cancer, colonoscopy is now a recommended procedure for patients over the age of 50 years as part of a broader screening program. The same tools are used to direct the biopsy and remove suspicious lesions. Challenges persist in the characterization of flat lesions and the introduction of optical imaging techniques beyond that of white light is a current area of research. Endoscopy is also a routine part of cancer management of many other sites including cancer of the head and neck, lung, and esophagus. The low cost and adaptability of endoscopes has resulted in other uses, including the design of radiation therapy treatments (Weersink et al, 2011) and the assessment of the response of tissue to cancer therapeutics. Georg et al (2009) have correlated rectal endoscopy scores with the volume of rectal tissue receiving elevated doses during radiation therapy of the cervix. The continued advances in endoscope technology and advances in optical imaging probes make this a particularly promising area of development for cancer imaging.