Instrumentation for Intraoperative Detection
Focusing on nuclear and optical modalities, the current chapter reviews the rapidly advancing technologies in intraoperative detection and imaging. This chapter reviews not only currently routine intraoperative imaging technologies but also investigational technologies potentially adaptable to the intraoperative setting.
KeywordsIntraoperative imaging Optical imaging Gamma probes Beta probes Cerenkov imaging Photoacoustic imaging Raman effect
Coherence anti-Stokes Raman scattering
- CdZnTe (or CZT)
Cadmium zinc telluride
Complementary metal oxide semiconductor
Sodium-doped cesium iodide
Thallium-doped cesium iodide
X-ray computed tomography
Diffuse optical tomography
Enhanced permeability and retention
Intraoperative fluorescence imaging
Gene encoding for firefly luciferase
Field of view
Green fluorescent protein
Cerium-doped gadolinium orthooxysilicate
Cerium-doped lutetium orthooxysilicate
Mega-Becquerel (106 Becquerel)
Magnetic resonance imaging
Multispectral optoacoustic tomography
Thallium-doped sodium iodide
Optical coherence tomography
Positron emission tomography
Peroperative compact imager
Tripeptide composed of L-arginine, glycine, and L-aspartic acid
Surface enhanced Raman scattering (or surface enhanced Raman spectroscopy)
Sentinel lymph node
Single-photon emission tomography
Small semiconductor gamma camera
Focusing on nuclear and optical modalities, the current chapter reviews the rapidly advancing technologies in intraoperative detection and imaging. This chapter reviews not only currently routine intraoperative imaging technologies but also investigational technologies potentially adaptable to the intraoperative setting. The reader is referred to a recently published volume entitled, “Imaging and Visualization in the Modern Operating Room: A Comprehensive Guide for Physicians,” which provides a detailed overview of the many logistical as well as technical considerations in modern intraoperative imaging .
Nuclear Counting and Imaging
Radionuclide-based detection and localization of tumors, especially small tumors, has several well-known limitations  which are mitigated through the use of intraoperative probes and, potentially, intraoperative gamma cameras. First, absolute tumor uptake of cancer-targeted radiotracers remains generally quite low, typically ~0.1% or less of the administered activity per gram. Second, overall radiation detection sensitivity in vivo is low as well, ranging from about 0.1% for gamma camera imaging (including SPECT) to ~10% for PET. This is exacerbated, of course, by the signal-degrading effect of attenuation of emitted radiation by overlying tissue. Third, a significant portion of the counts apparently emanating from a tumor or other targeted tissue may actually include counts originating elsewhere (i.e., from background activity in adjacent tissues) because of contrast- and resolution-degrading Compton scatter. However, because of the close proximity of a collimated detector to a tumor or sentinel lymph node which can be achieved at surgery, radionuclide detection of such structures can be enhanced using intraoperative probes or gamma cameras. In a study of simulated tumors in a torso phantom having uniform background activity, for example, Barber et al. demonstrated that a scintillation probe detected tumors with greater sensitivity than a gamma camera over a wide range of conditions provided the probe was placed within 1 cm of the tumor . Alternatively, tumor or sentinel lymph node detection may also be improved under some circumstances using beta (negatron or positron), rather than gamma, detection [9, 10, 11] because the very short range (typically ~1 mm or less) of such particulate radiations eliminates the contribution of confounding counts from activity other than in the immediate vicinity of the detector. Of course, the short range of particulate radiations also limits the application of beta probes to intraoperative or endoscopic settings with the lesion near the surface of the exposed tissue.
Beta probes. Because x- and γ-rays penetrate relatively long distances (circa 10 cm) of soft tissue, a major limitation of the use of gamma probes to specifically identify the target tissue in radioguided surgery is the presence of variable, generally high levels of background activity in normal tissues. Thus, even with a gamma probe centered over a tumor, the contribution of counts originating from activity in normal tissue underlying the tumor and even outside the field of view (due to penetration of the collimation and shielding) may degrade the tumor-to-normal tissue contrast (e.g., reducing tumor-to-normal tissue count ratios to less than 1.5:1 [15, 20, 21]) and thus tumor detectability to the point where lesions may be missed. A potential solution to this limitation of radioguided surgery is the use of so-called “beta” probes, that is, intraoperative probes which specifically count only charged particle (negatrons or positrons) radiation. Because they have such short ranges in soft tissues (typically of the order of 1 mm or less), beta particles emitted by a tracer source outside the probe’s FOV or underlying the surface tissue do not reach the detector and are not counted (by the same token, minimal if any collimation and shielding is required (Fig. 2c). As a result, the discrimination between higher-activity tumor and lower-activity normal tissues is enhanced (i.e., the tumor-to-normal tissue count ratios are increased). Of course, the short path length of beta particle restricts the use of such probes to surface lesions; beta probes could not be used, for example, for (percutaneous) detection of sentinel lymph nodes.
Intraoperative gamma cameras. The sensitivity and specificity of detection of sentinel lymph nodes using current approaches such as the preoperative gamma camera imaging, gamma probes, and “blue dye” technique are quite high. Newman , for example, performed a meta-analysis of nearly 70 published studies and found an overall sensitivity of over 90% and a false-negative rate of only 8.4% for detection of such nodes in breast cancer. For preoperative gamma camera imaging, detection rates of 72–85% have been reported . Sentinel lymph nodes were successfully detected using intraoperative gamma probes in 98% of patients in whom such nodes were successfully imaged preoperatively, with a false-negative rate of only 7%. And, for sentinel nodes not visualized by preoperative lymphoscintigraphy, there was a 90% detection rate intraoperatively. Importantly, however, negative preoperative lymphoscintigraphy often predicted a negative intraoperative probe result, and the foregoing improvement in the detection rate intraoperatively was primarily due to the use of blue dye . The American Society of Breast Surgeons has recommended a sensitivity of at least 85% and a false-negative rate of less than 5% as acceptable for sentinel node detection in breast cancer . There remains a need, therefore, to develop techniques to improve the sensitivity and the false-negative rates of sentinel lymph node detection. Intraoperative gamma camera imaging may provide the improvement required to satisfy the foregoing requirements. Mathelin et al. [27, 28, 29], for example, found that the use of an intraoperative small (5 × 5 cm) FOV gamma camera for detection of sentinel lymph nodes in breast cancer was practical. In a case report , intraoperative gamma camera imaging allowed detection of an additional sentinel lymph node (metastatic and with low radiotracer uptake) that was not detected by preoperative imaging or with a gamma probe, suggesting that intraoperative gamma camera imaging may reduce the false-negative rate. In addition to gamma camera and probe technology, the tracer employed (e.g., sulfur colloid versus nanocolloid, versus tilmanocept) and the injection site (e.g., intradermal versus subcutaneous, versus peri-tumoral) may account for the variation in lymph node detection.
Despite such promising preliminary data, it is not clear that the development and deployment of intraoperative gamma camera technology and the incremental improvement in the sentinel lymph node detection rate that such technology may provide will prove to be cost-effective. Certain considerations, however, lend support to the development of this technology. One such consideration is the variable level of proficiency among surgeons in gamma probe-based detection of sentinel lymph nodes : even with considerable training and experience, not all surgeons achieve a detection rate of 90% or better. In addition, certain sentinel lymph nodes are problematic anatomically or otherwise in terms of detectability. These include nodes which are unusually deep, close to (less than 30 mm from) the injection site or high-activity normal tissues, or have a low (less than 1%) radiotracer uptake [25, 29, 30, 31]. A gamma camera system having a spatial resolution of 3 mm or better at a distance (depth) of the order of 1 cm would likely visualize such problematic nodes intraoperatively. Such an imaging system would offer other practical advantages over probes: the signal is provided in the familiar format of a scintigraphic image rather than a numerical display or variable-frequency tone; the larger FOV of even small gamma cameras (several centimeters) than that of probes (less than 1 cm) allows more rapid interrogation of large areas and/or longer sampling, with collection of more counts and reduction in statistical uncertainty (noise); more straightforward reexamination of the surgical site post-lymph node excision to verify removal of foci of activity; and less reliance on potentially obliterated and otherwise ambiguous preoperative skin markings directing where measurements are to be performed intraoperatively . Intraoperative gamma camera systems thus merit development and evaluation.
A number of small FOV intraoperative gamma camera systems have been developed [12, 32, 33]. The earliest systems were handheld devices having FOVs of only 1.5–2.5 cm in diameter and using conventional NaI(Tl) or CsI(Tl) scintillation detectors. Later units used two-dimensional arrays (mosaics) of scintillation crystals connected to a position-sensitive PMT and, more recently, semiconductors such as CdTe or CdZnTe (CZT). The main problems with these early units were their very small fields of view and the resulting large number of images required to interrogate the surgical field and the difficulty in holding the device sufficiently still for the duration (up to 1 min) of the image acquisition. More recently, larger field-of-view devices have developed which are attached to an articulating arm for convenient and stable positioning. These systems are nonetheless fully portable and small enough overall to be accommodated in typical surgical suites.
Another semiconductor gamma camera, utilizing CdTe, was developed by Tsuchimochi and colleagues in Japan [36, 37, 38]. Their choice of CdTe was based on its superior uniformity (integral uniformity, 4.5%) and energy resolution (7.8%) compared to CZT. The camera, referred to as the “small semiconductor gamma camera (SSGC) ,” uses an array of 32 × 32 5-mm-thick CdTe elements, with a matrix of 1.2 × 1.2-mm pixels and a 4.5 × 4.5-cm FOV. The collimation, comprised of tungsten, had 1.2 × 1.2-mm square apertures to match the pixel arrangement. Spatial resolution without scatter was 3.9-, 6.3-, and 11.2-mm FWHM at 2.5, 5, and 10 cm, respectively. The 99mTc sensitivity at the surface without scatter was 300 cps/MBq, comparable to that of the POCI and better than that of a conventional gamma camera with LEHR collimation (~100 cps/MBq). The results of preliminary phantom and clinical imaging studies with the SSGC were encouraging.
The Institut Pluridisciplinaire Hubert Curien (Strasbourg, France) developed an intraoperative gamma camera known as the “CarollReS ” [27, 28, 29]. This device has a relatively large-area 50 × 50-mm cerium-doped gadolinium orthooxysilicate (GSO) scintillation crystal and parallel-hole collimation with 2-mm-wide apertures. Its 99mTc spatial resolution was 10-mm FWHM at 5 cm, sensitivity 130 cpm/kBq, and energy resolution 45%. A prototype version of this device with a larger 100 × 100-mm FOV has been fabricated as well. In a preliminary clinical study with the CarollReS camera, Mathelin et al.  compared the depth of lymph nodes estimated by imaging to their actual depth measured at surgery and found a generally good correlation, except in instances where only a portion of the sentinel lymph node was in the camera’s FOV. For 7 of 11 nodes whose depth could be estimated, the image-derived depth was correct.
Crystal Photonics GmbH (Berlin, Germany) markets a light (800-g) handheld gamma camera with a 5-mm-thick CZT detector having a 40 × 40-mm field of view, a photon energy range of 40–250 keV, and interchangeable tungsten or lead parallel-hole collimators. The spatial resolution at 3.5 cm is 5.4 mm with the high-resolution collimator to 9.2 mm with the high-sensitivity collimator, with the latter providing ~fourfold higher sensitivity than the former.
The SurgicEye GmbH (Munich, Germany) has introduced an intraoperative SPECT imaging technology termed “freehand SPECT” [45, 46]. In contrast to conventional SPECT systems, the Declipse® SPECT does not employ a gantry-mounted gamma camera rotating around the patient, but a handheld gamma probe interfaced to an infrared 3D tracking system. A 1- to 2-min scan is performed at anterior-posterior and lateral directions around the patient. The device can be used with low-energy photon emitters such as 99mTc and is compatible with different commercially available gamma probes. The Declipse™ SPECT website reports a reconstructed spatial resolution of 5 mm; details of the data acquisition and processing yielding this resolution were not specified, however.
Overall, small FOV gamma cameras have demonstrated detection rates for sentinel lymph nodes equal to or better than those of non-imaging gamma probes, despite having sensitivities (e.g., expressed in cps/MBq) typically about tenfold lower than those of such probes. The ability of such devices to image a surgical field intraoperatively and thus insure complete excision of lesions is a potentially useful enhancement of surgical management of cancer. The acquisition times per image are typically well under 1 min, so the overall duration of the surgical procedure should not be significantly prolonged. In addition, the use of pinhole collimation, despite having lower sensitivity than parallel-hole collimation, permits initial imaging at longer distance to visualize a larger anatomic area of interest followed by imaging at shorter distance to pinpoint and otherwise characterize suspicious foci of activity. Importantly, the scintigraphic image format is familiar to surgeons, likely facilitating clinical acceptance and integration of intraoperative imaging.
Intraoperative PET scanners . Prescient Imaging LLC is currently developing a portable compact whole-body PET scanner adaptable to intraoperative imaging . It incorporates a 360-degree detector with an axial FOV of 22 cm. The detectors are assembled into three sets, one planar and the other two circular arcs of 90° each connected with a hinge. One of the arcs is fixed while the other can be rotated about the hinge and thereby opened and closed. The planar detector is fixed horizontally such that it can fit under a patient table. By placing the moveable arc in the closed position, the 360° of detector coverage about the patient is achieved. Attenuation correction will be performed using a rotating rod transmission source.
Optical and Near-Infrared (NIR) Imaging
Because light emitted at any depth of tissue is scattered and otherwise dispersed as it passes through overlying tissue before emanating from the surface of the animal, the apparent size of the light source (Fig. 8c, d) is considerably larger than its actual size. Despite the excellent spatial resolution of the CCDs themselves, the effective resolution of optical and NIR imaging is generally rather coarse. Further, for planar optical and NIR imaging, the resulting images are only semiquantitative: absorption and scatter of the emitted light as it passes through overlying tissue makes the measured signal highly depth dependent. Thus, a focus of cells lying deep within tissue may appear less luminescent or fluorescent than an identical focus of cells at a more shallow depth; if excessively deep, such a focus of cells may be undetectable altogether. NIR radiations, however, have a substantially higher penetrability through tissue than blue to green radiations. Importantly, therefore, by using laser transillumination for excitation of administered NIR molecular probes in situ, tomographic fluorescence images can be mathematically reconstructed . The resulting three-dimensional images – in contrast to planar images – are at least semiquantitative: the signal intensity thus reconstructed is related to the local concentration of the fluorophore.
Bioluminescence imaging, because it requires genetic modification of the cells to be imaged, likely has very limited applicability in patients but has proven invaluable in preclinical research. However, with recent advances in adoptive immunotherapy of cancer, bioluminescence imaging conceivably may have some clinical utility in an intraoperative or endoscopic setting to assess tumor targeting of immune effector cells. To date, however, no such studies have been performed. Intraoperative, endoscopic, and even surface fluorescence imaging of patients has been performed and continues to advance (see below).
Narrowband imaging (NBI) is a notable refinement of non-fluorescence optical cystoscopy that improves the visualization of blood vessels and bladder mucosa and thereby enhances the contrast between cancerous and normal bladder epithelium (given that urothelial lesions are typically hypervascularized due to elevated microvessel density) . NBI exploits these angiogenic features of bladder cancers by filtering white light into two discrete wavelength bands, one blue (415 nm) and one green (540 nm), both of which are absorbed by hemoglobin. The shorter-wavelength 415-nm light penetrates only the superficial layers of the mucosa and yields brownish images of the superficial capillaries. The longer-wavelength 540-nm light penetrates deeper into the bladder wall and produces greenish images. Bladder tumors are thus visually identified by the intensity of brown-green coloration that is characteristic of the elevated vessel density of tumors and distinct from the pink-to-white coloration of the normal mucosa.
Cerenkov imaging. Cerenkov imaging is a new approach to optical imaging based on the emission of a continuum of visible light associated with the decay of certain radionuclides (actually, with the particles emitted as result of the radionuclide decay) [64, 65, 66, 67, 68, 69, 70, 71, 72, 73]. This phenomenon, now known as the “Cerenkov effect,” was first observed in the 1920s and characterized in the 1930s by Pavel Cherenkov . In 1958, Cerenkov shared the Nobel Prize in Physics with colleagues Ilya Frank and Igor Tamm for the discovery and explanation of the effect which now bears his name. Cerenkov radiation is perhaps familiar to some readers as the bluish “glow” observed in the water pools containing spent, but still radioactive, fuel rods at nuclear reactors. It arises when charged particles such as beta particles travel through an optically transparent, insulating medium at a speed greater than that of light in that medium. The Cerenkov effect, often analogized to the sonic boom that occurs at the instant a supersonic plane exceeds the speed of sound in air, occurs as the charged particles dissipate their kinetic energy by polarizing the electrons in the insulating medium (most commonly, water) as they travel through the medium. As these polarized electrons then relax (or re-equilibrate), and if the charged particle is traveling faster than light, constructive interference of the light thus emitted occurs, producing the grossly visible Cerenkov radiation.
Photoacoustic imaging. In photoacoustic imaging [76, 77, 78], tissues are illuminated with laser light. When radiofrequency (RF) pulses are used, the technology is termed, “thermoacoustic imaging.” Some of the delivered energy will be absorbed and converted into heat, leading to transient thermoelastic expansion of the illuminated tissue and thus ultrasonic (i.e., MHz-frequency) emissions. The ultrasonic waves thus emitted are then detected by ultrasonic transducers to form images. Image contrast is provided by the differential absorption among tissues of the incident excitation light. In contrast to fluorescence imaging, in which scattering in tissue degrades spatial resolution with increasing depth, photoacoustic imaging provides better spatial resolution (of the order of 100 μm) and deeper imaging depth (of the order of 1 cm or greater) because there is far less absorption and scattering in tissue of the ultrasonic signal compared to the emitted light signal in fluorescence imaging. When compared with ultrasound imaging, in which the contrast is limited because of the similarity in acoustical properties among tissues, photoacoustic imaging provides better tissue contrast as a result of the wider range of tissue optical properties. The optical absorption in biological tissues can be due to endogenous molecules such as hemoglobin or melanin or exogenously administered contrast agents. Since blood exhibits orders of magnitude higher light absorption than other tissues, there is sufficient endogenous contrast provided by oxygenated hemoglobin (HbO2) and deoxygenated hemoglobin (Hb) for photoacoustic imaging to visualize blood vessels.
Photoacoustic imaging has been used successfully in preclinical models for tumor perfusion angiogenesis monitoring (see Fig. 19 ), blood oxygenation mapping, functional brain imaging, and melanoma detection, among other applications. The resulting functional images can be superimposed on high-resolution B-mode anatomic images.
Diffuse optical tomography. Diffuse optical tomography (DOT) utilizes NIR light to generate quantitative functional images of tissue with a spatial resolution of 1–5 mm at depths up to several centimeters [83, 84]. Propagation of NIR light through a medium is dominated by scattering rather than absorption – tissue absorption path lengths are ~10 cm while scattering path lengths are less than 50 μm – and can be modeled as a diffusion process where photons behave stochastically (in a manner analogous to that of particles in random-walk modeling of diffusion). Quantitative measurements can be obtained by separating light absorption from scattering using spatial or temporal modulation techniques. Tissue molecular composition, including the determination of the concentrations of oxy- and deoxyhemoglobin, water, lipid, and exogenous probes, and tissue structure can be determined from absorption and scattering measurements, respectively. Time modulation systems use picosecond optical pulses and time-gated photon-counting detectors; frequency modulation systems use a RF-modulated light source, PMTs or fast photodiodes, and RF phase detectors. DOT has been applied to breast cancer diagnostics, joint imaging, and blood oximetry (i.e., activation studies) in the human muscle and brain tissue as well as to cerebral ischemia and cancer studies in small animals and is adaptable to intraoperative imaging. Commercial instruments are now available that yield tomographic and volumetric image sets. These devices are compact, portable, and relatively inexpensive (~$150 K).
Optical coherence tomography. Optical coherence tomography (OCT) is an interferometric technique, typically employing low-coherence NIR light, to produce two-dimensional images of tissue surface layers and structure . The principle of OCT is analogous to that of pulse-echo (i.e., B-mode) ultrasound imaging except OCT uses light instead of acoustic waves to delineate tissue structure by measuring reflectance of light rather than sound waves and thus achieves far better spatial resolution but with less depth penetration. The technique has been described as “an optical biopsy,” since OCT can produce near-histologic images (spatial resolution, 1–15 μm) without excision. Due to photon absorption and scattering, its sampled depth is limited to within several millimeters of the tissue surface. The two-dimensional images can be assembled to construct a volumetric image set. In OCT, the axial resolution is proportional to the center wavelength and inversely proportional to the bandwidth of the light source and improves with the index of refraction of the sample. Originally developed for and still most commonly applied to ophthalmology (to obtain detailed images of retinal structure), OCT is being applied for cancer diagnosis, evaluation of coronary artery disease and tissue characterization.
The last decade has featured remarkable technical advances in intraoperative imaging and techniques potentially adaptable to intraoperative imaging. While intraoperative nuclear counting using handheld probes is now routine, intraoperative nuclear imaging remains largely investigational. While certain applications of intraoperative and endoscopic optical imaging (such as optical cystoscopy) are also now fairly routine, more exotic forms of optical imaging such as Cerenkov imaging and photoacoustic imaging have only recently been translated to the clinic. Other promising forms of optical imaging such as Raman spectroscopic imaging have not yet been clinically translated to any significant extent. Multimodality intraoperative probes capable of simultaneously detecting both nuclear and optical signals are being developed as well. The ultimate clinical impact of these new imaging modalities in intraoperative and endoscopic settings remains to be determined.
- 3.Cody III HS, editor. Sentinel lymph node biopsy. London: Martin Dunitz; 2002.Google Scholar
- 4.Mariani G, Giuliano AE, Strauss HW, editors. Radioguided surgery: a comprehensive team approach. New York: Springer; 2008.Google Scholar
- 6.Gulec SA, Moffat FL, Carroll RG. The expanding clinical role for intraoperative gamma probes. In: Freeman LM, editor. Nuclear medicine annual 1997. Philadelphia: Lippincott-Raven Publishers; 1997. p. 209–37.Google Scholar
- 7.Woolfenden JM, Barber HB. Intraoperative probes. In: Wagner HN, Szabo Z, Buchanan JW, editors. Principles of nuclear medicine. 2nd ed. Philadelphia: WB Saunders; 1995. p. 292–7.Google Scholar
- 13.Zanzonico P. The intraoperative gamma probe: design, safety, and operation. In: Cody III HS, editor. Sentinel lymph node biopsy. London: Martin Dunitz; 2008. p. 45–68.Google Scholar
- 24.Newman LA. Current issues in the surgical management of breast cancer: a review of abstracts from the 2002 San Antonio Breast Cancer Symposium, the 2003 Society of Surgical Oncology annual meeting, and the 2003 American Society of Clinical Oncology meeting. Breast J. 2004;10 Suppl 1:S22–5.CrossRefPubMedGoogle Scholar
- 62.Bradbury M, Pauliah M, Wiesner U. Ultrasmall fluorescent silica nanoparticles as intraoperative imaging tools for cancer diagnosis and treatment. In: Fong Y et al., editors. Imaging and visualization in the modern operating room: a comprehensive guide for physicians. New York: Springer; 2015. p. 167–79.CrossRefGoogle Scholar
- 63.Bradbury MS, Pauliah M, Zanzonico P, et al. Intraoperative mapping of sentinel lymph node metastases using a clinically translated ultrasmall silica nanoparticle. Wiley Interdiscip Rev Nanomed Nanobiotechnol. 2015.Google Scholar
- 65.Cerenkov PA. Visible emission of clean liquids by action of gamma-radiation. C R Dokl Akad Nauk SSSR. 1934;2:451–4.Google Scholar
- 76.Xu MH, Wang LHV. Photoacoustic imaging in biomedicine. Rev Sci Instrum. 2006;77:41–101.Google Scholar