Skin cancer is the most common form of malignancy, with more than one million cases reported in the United States every year [1
]. A majority of these cases are nonmelanoma skin cancers that are usually not life threatening and can be treated. For melanoma skin cancer, early detection is critical to the survival of the patient. At advanced stages, melanomas metastasize and spread to other organs, significantly decreasing five-year survival (~15%) [1
]. The current procedure for the detection of skin cancers is a clinical examination by a qualified dermatologist or general practitioner followed by a tissue biopsy and histopathology—the current gold standard—to confirm the results. However, there are several limitations associated with the process. Tissue biopsies are time-consuming, invasive, and subjective procedures that have led to an increase in the morbidity associated with the diagnosis of skin cancers. Additionally, a recent clinical study showed that approximately 40% of suspicious biopsied lesions were classified as benign by pathology [2
]. Therefore, there is a need for a real-time, noninvasive diagnostic method that can reduce the morbidity and mortality associated with skin cancers.
Optical techniques such as diffuse optical spectroscopy and laser-induced fluorescence spectroscopy offer a noninvasive alternative to tissue biopsies for determining the state of tissue. Light is delivered and collected with optical fiber probes that are placed in contact with the skin surface. The weak light pulses sample the tissue beneath the probe noninvasively and provide information regarding tissue morphology, function, and biochemical composition. As these physiological parameters change with progression of disease, optical spectroscopy offers a means to measure disease progression without tissue removal.
A number of studies have demonstrated the feasibility of diffuse reflectance spectroscopy (DRS) for the optical biopsy of sampled tissue using either a model-based [3
] or empirical approach [4
]. The model-based analysis of diffuse reflectance has the advantage of providing quantitative measures of the wavelength-dependent reduced scattering (μs
′) and absorption coefficients (μa
) that relate to tissue morphology and function, respectively. The reduced scattering coefficient in tissue can be described as an inverse power law function of wavelength in the visible range and is a function of the scatterer size and density [6
]. The absorption coefficient of tissue is a function of physiological parameters such as blood volume fraction, oxygen saturation, blood vessel diameter, and melanin concentration. Each parameter provides a different perspective of sampled tissue, and a number of them have been demonstrated to change with progression to dysplasia in various organs [3
Laser-induced fluorescence spectroscopy (LIFS) has been widely used to detect dysplasia in various organs [9
]. LIFS is based on the study of prominent tissue fluorophores such as nicotinamide adenine dinucleotide (NADH), collagen, and flavin adenine dinucleotide (FAD). NADH and FAD are important indicators of metabolic activity. Hence, large-scale cell proliferation or tumor growth can be identified by significant changes in NADH and FAD fluorescence [13
]. In addition, studies have reported an increase in production of matrix metalloproteinases (MMPs) with progression to dysplasia [15
]. Metalloproteinases are a group of collagenases that cleave the collagen cross links, leading to a reduction in the levels of collagen fluorescence. Coghlan et al.
demonstrated that such spectral changes in fluorescence precede morphological changes in tissue, thus offering a means for the early detection of dysplasia [16
Although DRS and LIFS are independently capable of measuring tissue pathology, a combination of both techniques can provide complementary information about tissue morphology, function, and biochemical composition. In addition, acquiring DRS in combination with LIFS allows for correcting the fluorescence spectra distorted by absorption and scattering events [17
]. Cottrell et al.
] developed a clinical instrument that used both reflectance and fluorescence spectroscopy for monitoring the optical properties of basal cell carcinoma during treatment with photodynamic therapy. Ramanujam and collaborators have conducted extensive studies on malignant breast tissue, using a combination of DRS and intrinsic fluorescence (corrected LIFS) spectroscopy (IFS) [20
]. Clinical studies in the cervix [23
] and oral cavity [24
] using combined DRS, IFS, and model-based light scattering spectroscopy have demonstrated that a multimodal approach provides a superior tool for differentiating between normal and dysplastic tissue compared with any one method alone. A combination of DRS and IFS has been used for in vivo
cancer diagnosis in the breast with improved results [25
The clinical studies described earlier use systems that collect white light reflectance and fluorescence excitation-emission matrices (EEMs) over a wide wavelength range. Zângaro et al.
] developed a clinical instrument (fast-EEM) for simultaneously measuring spectrally resolved reflectance and fluorescence excitation-emission matrices (EEMs) in the wavelength range 350–700 nm that was subsequently used in diagnostic studies of the cervix [23
], oral cavity [24
], and esophagus [27
]. The instrument consisted of an independent white light source, a nitrogen laser that pumped a series of 10 dye cuvettes to provide 11 different excitation wavelengths (337–505 nm) for fluorescence, and a multichannel diode detector for collecting tissue spectra. Subsequently Müller et al.
] and Tunnell et al.
] developed upgraded versions of the fast-EEM instrument that extended the wavelength range of collection (300–800 nm) and significantly improved the data acquisition time to less than 1 s. The improved fast-EEM was used in studies in the oral cavity [30
], breast [25
], and atherosclerotic plaques [31
]. Zuluaga et al.
] developed the FastEEM instrument for collecting fluorescence EEMs and spectrally resolved diffuse reflectance spectra at multiple source–detector separations. This instrument utilized a single light source coupled with filters for delivering white light (380–950 nm) and 18 fluorescence excitation wavelengths (330–500 nm) and was subsequently used in clinical studies to detect dysplasia in the cervix [33
]. Freeberg et al.
] have reported upgraded versions of this FastEEM instrument that was used at multiple sites for studies in the cervix [36
]. Ramanujam and collaborators used both custom-built [37
] and commercially available [22
] clinical instruments that collected spectrally resolved diffuse reflectance at three different source–detector separations and fluorescence EEMs for nine excitation wavelengths.
We have developed a compact, portable and fast clinical instrument that combines DRS and IFS for the early detection of melanoma and nonmelanoma skin cancers. We measure white light reflectance in the wavelength range of 350–700 nm. However, in contrast to the clinical systems described previously, we employ only two excitation wavelengths for fluorescence measurements: 337 nm for exciting NADH and collagen and 445 nm for FAD. Several studies have identified the fluorescence from these excitation bands to be diagnostically relevant for separating normal and dysplastic tissue [25
]. Using only two excitation wavelengths has two important advantages. First, it facilitates simple and compact instrument design by eliminating the need for large excitation and emission monochromators for collecting fluorescence EEMs. This makes the system portable and clinically compatible. Second, we can significantly reduce data acquisition and processing time because we do not record full fluorescence EEMs. We use a microelectromechanical systems-based (MEMS-based) fiber optic switch for controlling the excitation sequence of the white light and laser sources, instead of a filter wheel. This reduces the instrument complexity and eliminates losses due to additional optics. We use an interline transfer CCD cooled to −30°C that allows data collection with short integration times (50 μ
s) while affording a high signal-to-noise ratio by eliminating thermal noise and dark current. Using such gated detection techniques allows us to leave the room lights on while recording data, an important step toward developing a clinically compatible system. We use a fiber optic probe with a close source–detector separation of 250 μ
m to ensure a superficial sampling depth, because skin cancers originate in the superficial epidermal layer of tissue. In this paper, we discuss the design and instrumentation of a combined DRS–IFS clinical system and detail the features that make it clinically compatible, portable, and capable of rapid data acquisition from tissue. We also briefly describe the algorithms used to extract both optical properties and fluorophore contributions from in vivo
measurements. The development and validation of a model to extract optical properties has already been discussed in a previous publication [40
]. We demonstrate the accuracy of the system on tissue-simulating phantoms over a wide range of optical properties. We also present in vivo
measurements from clinically normal skin, dysplastic nevi, and malignant non-melanoma skin cancers.