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The detection of a gold nanoparticle contrast agent is demonstrated using a photothermal modulation technique and phase sensitive optical coherence tomography (OCT). A focused beam from a laser diode at 808 nm is modulated at frequencies of 500 Hz–60 kHz while irradiating a solution containing nanoshells. Because the nanoshells are designed to have a high absorption coefficient at 808 nm, the laser beam induces small-scale localized temperature oscillations at the modulation frequency. These temperature oscillations result in optical path length changes that are detected by a phase-sensitive, swept source OCT system. The OCT system uses a double-buffered Fourier domain mode locked (FDML) laser operating at a center wavelength of 1315 nm and a sweep rate of 240 kHz. High contrast is observed between phantoms containing nanoshells and phantoms without nanoshells. This technique represents a new method for detecting gold nanoparticle contrast agents with excellent signal-to-noise performance at high speeds using OCT.
©2008 Optical Society of America
1. Introduction
Optical coherence tomography (OCT) is a high-resolution biomedical imaging modality that produces cross-sectional and three-dimensional images of tissue microstructure by interferometrically measuring the amplitude and echo time delay of backscattered light [1]. OCT imaging can derive contrast from sources that are either endogenous or exogenous to the tissue being imaged. The most utilized source of endogenous contrast is spatial variations in the scattering properties of the tissue, which produces contrast in conventional OCT images. In this case, only the amplitude of the interference signal is analyzed to form the image. Another source of endogenous contrast is velocity or flow. Typically referred to as Doppler OCT or optical Doppler tomography (ODT), these techniques analyze phase changes in the interference signal over brief time periods to detect vascular blood flow [2–7]. Endogenous OCT contrast can also be derived from variations in the size of scattering particles in the tissue [8–10] or wavelength-dependant absorption of different tissue components [11, 12] using spectroscopic OCT [13–15]. Finally, non-centrosymmetric endogenous tissue components, such as collagen, can be detected using second harmonic OCT [16–21].
Exogenous contrast agents have not typically been used in OCT, but recently have been studied more closely. As in other biomedical imaging modalities, OCT contrast agents promise to enable enhanced visualization of selected features such as microvasculature, epithelial structures, and diseased or abnormal tissue. Agents such as methylene blue, rhodamine, and indocyanine green can be detected by signatures in their electron relaxation times using pump-probe OCT [22–24]. OCT contrast enhancement has also been demonstrated using scattering microspheres [25], near-infrared (NIR) dyes [26], iron oxide microparticles [27], and, more recently, nanoparticles [28–36].
Gold nanoparticles are particularly attractive contrast agents since they can be targeted to biochemical markers associated with specific types of disease such as cancer [37, 38], which suggests the possibility of highly sensitive and specific OCT detection of early neoplasia. Gold nanoshells consist of an inner silica core surrounded by a thin gold shell. By changing the relative dimensions of the core and shell, the optical resonance frequency of the particles can be tuned from ultraviolet to near infrared wavelengths [28]. This allows customized tailoring of the optical scattering and absorption properties of the particles to suit the needs of the specific application. Gold nanoshells are highly biocompatible, water-soluble, and commercially available. Nanoshells can be designed with high absorption, targeted to cancer cells, and used for photothermal therapy with minimal damage to surrounding tissue [39]. Other types of gold nanoparticles such as nanorods [33] and nanocages [29] exhibit similar properties to nanoshells and can also be used for exogenous OCT contrast enhancement.
Methods for detecting exogenous contrast agents using OCT can be divided into two general categories: passive techniques and active techniques. Passive detection techniques rely on time-invariant differences in the optical properties of the agent compared to the tissue to generate contrast. Differences in the absorption of near infrared dyes compared to tissue can be passively detected using spectroscopic OCT [26]. Similar spectroscopic methods have been applied to detect gold nanoparticles, where the absorption of the nanoparticles caused a blue shift of the OCT signal [29]. Optical scattering can also be used to detect gold nanoparticles using conventional amplitude-based OCT [28, 30, 32–34, 36] since the peak scattering or absorption wavelength of the nanoparticles can be selected to overlap with the OCT imaging wavelength. Passive contrast agent detection may be difficult to apply in vivo, however, since the signal is not background-free and variations in the optical properties of heterogeneous tissue can mask the scattering and absorption characteristics of the agents.
Active contrast agent detection techniques modulate a property of the agent to enhance visualization against a heterogeneous tissue background. One example of active contrast agent detection is magnetomotive OCT [27, 31, 35]. In this technique, superparamagnetic iron oxide (SPIO) nanoparticles are taken up by cells in the sample tissue and are then exposed to an external magnetic field of 0.06–0.5 T that is modulated at 3–50 Hz. Modulation of the external magnetic field causes localized motion in regions of the tissue that have taken up the SPIO, and this motion is detected by fluctuations in the amplitude [31] or phase [35] of the OCT interference signal. Magnetomotive OCT achieves a high signal-to-noise ratio (SNR) for detecting SPIO nanoparticles since active modulation of the contrast agent results in a detection scheme that is less susceptible to background noise. However, this technique requires the application of fairly strong magnetic field gradients (up to 11 T/m) and is limited in imaging speed due to the relatively slow mechanical response of SPIO-laden tissue. These factors may make magnetomotive OCT challenging to apply for in vivo imaging in humans.
Here we demonstrate an active contrast agent detection technique for high-speed OCT imaging based on photothermal modulation. The technique uses gold nanoshells designed to have high absorption at 808 nm where tissue absorption is inherently low. A multimode laser diode operating at 808 nm is used to induce small-scale, localized temperature gradients in regions of the sample that contain the contrast agent. These temperature variations alter the optical path length in the sample. Changes in path length are detected using a swept source / Fourier domain OCT phase microscopy system [40–43] built using a double-buffered Fourier domain mode locked (FDML) laser operating at 1315 nm and a sweep rate of 240,000 sweeps per second (240 kHz). By modulating the 808 nm laser diode at a known frequency and observing variations in optical path length that occur only at that frequency, the contrast agent can be detected in a way that significantly reduces background noise. Contrast agent SNR’s of up to 131 are obtained using modulation frequencies of 500 Hz–60 kHz.
This technique is similar in some ways to photothermal lensing methods that have been previously applied for detecting metal nanoparticles [44, 45]. In photothermal lensing, nanoshell-laden samples are exposed to a pump laser that causes localized heating and alters the diffractive properties of the sample. A probe beam at a different wavelength is defocused or diffracted by the heated region, and this effect is observed in transmission in the far field. The method described here, however, relies on detecting changes in optical path length instead of a diffractive or defocusing effect. By detecting path length changes with phase-sensitive OCT, nanometer-scale displacements can be measured. Additionally, depth selectivity is obtained by the use of low-coherence interferometry in the form of OCT. Further enhanced SNR’s are achieved by high-speed modulation of the thermal excitation beam. The technique described here can be integrated with 3D-OCT imaging to provide contrast-enhanced images of tissue architectural morphology. In the future, photothermal detection of gold nanoshells using high-speed, phase-sensitive OCT may enable targeted in vivo imaging of disease with high sensitivity and specificity.
2. Experimental setup
2.1 Double-buffered FDML laser
The phase-sensitive OCT system uses a new type of FDML laser. FDML lasers are a class of wavelength-swept light source that overcome fundamental limitations in sweep speed present in conventional swept lasers [46, 47]. In FDML lasers, the intracavity filter element is tuned synchronously with the optical roundtrip time of the light in the cavity. This results in a quasi-stationary operating regime where the tunable filter is at the same position every time a given wavelength in the sweep returns to the filter. FDML lasers have previously been demonstrated with tuning ranges of up to 160 nm at a center wavelength of 1315 nm, average output powers of up to 35 mW, and sweep rates of up to 370 kHz [48, 49]
The first generation of FDML lasers used fiberoptic Fabry-Perot tunable filters (FFP-TF’s) driven with sinusoidal waveforms and semiconductor optical amplifiers (SOA’s) driven with a continuous-wave signal. This generated an output that was an alternating series of short-to-long (forward) and long-to-short (backward) wavelength sweeps. At high sweep rates or long ranging depths, however, the noise performance of the forward sweep degrades more rapidly than the backward sweep. To avoid this problem, the second generation of FDML lasers used a “buffered” configuration to create unidirectional sweeps with low noise levels [48]. In buffered FDML lasers, time-shifted copies of the sweep are extracted from evenly-spaced points in the cavity. During the time normally occupied by the undesired forward sweep, the cavity SOA is modulated off. The copies of the remaining backward sweep are recombined in an external fiber coupler, enabling unidirectional sweeping without decreasing the duty cycle.
In the current system, the concept of buffered FDML is extended to further multiply the sweep rate and increase the sweep linearity. The light source used for these experiments is a “double-buffered” FDML laser, where two copies of the backward sweep are extracted from the cavity and then routed to a second external buffering stage. External buffering was recently demonstrated in an FDML laser operating at a center wavelength near 1050 nm for ophthalmic imaging [50]. In the external stage, four copies of the original sweep are created, time-shifted, and recombined in a final fiberoptic coupler. This arrangement quadruples the effective sweep rate compared to the FFP-TF drive frequency. The use of external buffering also minimizes power losses during recombination of the sweep copies. This concept can be extended by adding additional stages composed of a 50/50 splitter and fiber delay line to the external buffer. Each additional stage can further multiply the sweep rate without adding additional loss beyond the excess loss of the splitter and propagation loss of the fiber. To prevent temporal overlap of the sweep copies, the FFP-TF is driven with a high amplitude to reduce the amount of time required for each sweep. The duty cycle of the intracavity SOA is also correspondingly reduced. Double buffering has the added benefit of decreasing the portion of the sine wave used to generate the sweep, which improves the linearity of the optical frequency sweep.
Figure 1 shows a schematic of the double-buffered FDML laser. The fiber cavity is 3.4 km long (SMF-28e, Corning Inc) and contains two optical isolators, a broadband SOA, and a FFP-TF (Lambda Quest LLC). The FFP-TF is driven with a 5.8 VAC sinusoidal waveform at 59.8 kHz. At this drive amplitude the FFP-TF tunes over ~1.8 free spectral ranges, each of 170 nm, in 8.4 µs. The SOA is modulated with a square wave of 4.5 µs, such that only the backward sweeps are generated, with a total tuning range of 158 nm at a center wavelength of 1315 nm. The duration of each backward sweep is 3.9 µs, slightly shorter than the duration of the SOA modulation window due to the finite frequency response of the SOA. Two copies of the backward sweep are extracted at evenly-spaced points within the cavity using 80/20 and 70/30 fiberoptic splitters. These copies are again split, copied, time-delayed by 4.1 µs, and recombined in the external buffering stage. The output of the external buffering stage is amplified by a second broadband SOA.
Fig. 1. Double-buffered Fourier domain mode locked (FDML) laser, operating at a sweep rate of 240 kHz with a tuning range of 158 nm at a center wavelength of 1315 nm. Sweep rate is quadrupled by internal and external buffering stages. ISO, optical isolator. SOA, semiconductor optical amplifier. FFP-TF, fiber Fabry-Perot tunable filter.
The final output of the double-buffered FDML laser is a train of wavelength sweeps with a sweep rate of 239.2 kHz, exactly four times the FFP-TF drive frequency. The time-averaged output spectrum, measured using an optical spectrum analyzer, is shown in Fig. 2(a). The total tuning range is 158 nm and the full width at half maximum (FWHM) is 117 nm. The average output power is 62 mW with a duty cycle of 91%. OCT point spread functions (PSF’s) measured at increasing ranging depths are shown in Fig. 2(b). The sensitivity decreases by 5.5 dB at a ranging depth of 2 mm in air and by 23 dB at a ranging depth of 6 mm.
Fig. 2. (a). Integrated FDML output spectrum, with a tuning range of 158 nm and a full-width-half-maximum bandwidth of 117 nm. (b). OCT point spread functions measured at increasing ranging depths.
In a double-buffered FDML laser, only 1/4 of the FFP-TF drive period is used to produce each sweep. In a single-buffered FDML laser, 1/2 of the period is used. Since a smaller portion of the sine wave is used to create the sweep in a double-buffered FDML laser, the FFP-TF tunes more linearly in k. This improves data acquisition efficiency by reducing unnecessary samples during slow portions of the sine wave. After k recalibration, the PSF from the double-buffered laser gives an axial resolution of 8.3 µm in air or 6.0 µm in tissue.
2.2 Swept source OCT phase microscope
The swept source OCT phase microscope is similar to the design previously described in reference [43], but has been modified to collinearly direct the 808 nm laser diode beam and the 1315 nm OCT beam onto the sample. Figure 3 shows a schematic of the system. 95% of the FDML output is routed to a common path interferometer, designated as the “sample interferometer.” The liquid sample is held in a glass cuvette, where the first glass/liquid interface provides the reference reflection for the interferometer. The output of the fiber-pigtailed 808 nm laser diode (Power Technology Inc) is combined with the OCT beam using a dichroic mirror. The diode has a maximum output power of 300 mW and is pigtailed to a multimode fiber with a 50 µm core diameter. The 808 nm beam is collimated by a 15 mm focal length lens and the 1315 nm OCT beam is collimated by a 20 mm focal length lens.
An XY pair of galvanometer mirrors with a 6 mm clear aperture is used to aim the combined 808 nm / 1315 nm beam on the sample. The beams are focused using a 30 mm focal length achromatic objective lens. The 808 nm beam diameter is ~140 µm at the 1/e2 intensity point, while the 1300 nm beam diameter is 15 µm at the 1/e2 intensity point. The 808 nm spot size was measured with a CCD camera, while the 1300 nm spot size was estimated using a resolution test target. The remaining 5% of the FDML laser output is routed to a second common path interferometer, designated as the “calibration interferometer,” which uses a 210 µm thick glass slide as the sample. The front and back air/glass boundaries generate two fields that interfere to produce a calibration signal for resampling the sample fringes onto a linear k spacing, and for removing slow phase drift caused by the FDML laser [43]. The sample and calibration data are acquired simultaneously using a 2 GS/s, 8 bit digital oscilloscope (Tektronix Corp), and processing is performed post-acquisition using a personal computer. The 808 nm laser diode is modulated by a digital pulse generator that is synchronized to the beginning of each wavelength sweep.
Fig. 3. Swept-source OCT phase microscope with photothermal modulation system. C1, C2, C3, collimating lenses. OBJ, objective lens. DCM, dichroic mirror. X,Y, galvanometer mirrors. PD, photodiode. A, amplifier. TRG, sweep trigger. CH 1, OCT signal input. CH 2, calibration signal input. DAQ, data acquisition. Inset shows measured phase noise of 2.2 mrad.
The phase noise of the system was measured by placing a 210 µm glass slide in the sample interferometer and recording the position of the back surface relative to the front surface over 30 ms. The slow component of the phase drift caused by the FDML laser was recorded using the calibration interferometer and subtracted from the sample data. The results of this measurement are shown in the inset of Fig. 3. The phase noise was measured to be 2.2 mrad, corresponding to a displacement sensitivity of 153 pm.
2.3 Sample preparation
Gold nanoshells with a 120 nm core, 16 nm shell, and peak absorption wavelength of 780 nm were obtained commercially (Nanospectra Biosciences Inc.). The nanoshells were mixed with deionized water and diluted to a concentration of 1010 mL-1. At this concentration, the absorption coefficient at 808 nm was approximately 3.88 cm-1 with a FWHM bandwidth of ~400 nm. A glass cuvette with a sample path of 200 µm was filled with the solution, and the cuvette was placed in the sample interferometer of the OCT phase microscope. As shown in Fig. 4, the glass/fluid interface between the cuvette cover and the nanoshell solution was used as the reference reflection for the common path interferometer. The fluid/glass interface between the nanoshell solution and the cuvette body was monitored for small changes in optical path length, corresponding to localized absorption and heating of the nanoshells during exposure to the 808 nm laser beam. For control experiments, pure deionized water was placed in the same cuvette instead of the gold nanoshell solution.
The 1315 nm OCT beam was focused to a 1/e2 width of ~15 µm by the objective lens in the sample interferometer. The 808 nm beam was focused to a 1/e2 width of ~140 µm by the same objective, due to the larger numerical aperture and core size of the multimode fiber attached to the laser diode. A smaller diameter is generally more desirable for the 808 nm beam in order to increase the energy density and induce larger optical path changes in the sample. However, the larger beam diameter ensured uniform heating in the volume interrogated with the OCT system and also simplified alignment of the 808 nm beam to the OCT beam.
Fig. 4. Sample holder and beam geometries for photothermal detection of gold nanoparticles. Beam widths are approximate 1/e2 points of optical intensity.
3. Results
3.1 Photothermal detection of gold nanoshells
To demonstrate the concept of photothermal detection of gold nanoshells, OCT phase microscopy was performed on both pure deionized water and a gold nanoshell solution, both with and without exposure to the 808 nm laser source. The results of these experiments are shown in Fig. 5. In Fig. 5, all plots show the interference fringe phase associated with the second fluid/glass interface, measured over time, at a single spatial location in the cuvette. Figure 5(a) shows the measured phase when the sample is pure deionized water and the 808 nm laser is disabled. The phase profile is generally featureless and corresponds to noise. Figure 5(b) shows the measured phase when the sample is a 1010 mL-1 gold nanoshell solution and the 808 nm laser is disabled. Phase noise is increased due to increased scattering in the sample, but no systematic pattern is observed. This nanoshell concentration is consistent with estimates of concentrations that may be attainable in tumor tissue following systemic administration of antibody-labeled nanoshells [32].
Figure 5(c) shows the measured phase from the deionized water sample, but with the 808 nm laser activated. The 808 nm laser was set to provide 276 mW and was modulated with a 500 Hz square wave with a 50% duty cycle, giving an average power of 138 mW. The 1310 nm FDML laser provided an additional 20 mW of power on the sample. The 808 nm laser modulation pattern is shown at the top of the plot, and the vertical line indicates the time at which the 808 nm laser was switched on at t=4 ms. No change in the phase is observed compared to Fig. 5(a), indicating that the absorption of water at 808 nm is not high enough to cause localized heating and induce optical path changes. Figure 5(d) shows the measured phase from the nanoshell solution with the 808 nm laser activated. The modulation parameters were identical to those used for the deionized water sample, and the vertical line indicates the time at which the laser was switched on at t=4 ms. In this case, a strong phase response is observed. The high absorption of the gold nanoshells at 808 nm causes localized heating of the solution, which in turn increases the optical path length of the sample. The phase response of the sample shows the same modulation pattern as the 808 nm laser. Each phase modulation Δϕ is ~1.1 rad peak-to-peak, corresponding to a physical path difference ΔL of ~87 nm using ΔL=λ 0Δϕ/4πn, where λ 0=1315 nm is the center wavelength of the FDML laser and n=1.33 is the refractive index of water. Since there is insufficient time for the solution to fully cool using these 808 nm modulation parameters, there is a slow increase in temperature producing a cumulative increase in optical path of ~7.5 rad or 590 nm over 30 ms.
Fig. 5. Measured phase from back surface of cuvette vs. time. a, Deionized water with 808 nm laser deactivated. b, 1×1010 mL-1 nanoshell solution with 808 nm laser deactivated. c, Deionized water with 808 nm laser modulated at 500 Hz. Red pulse train shows laser modulation signal. d, 1×1010 mL-1 nanoshell solution with 808 nm laser modulated at 500 Hz. Red pulse train shows laser modulation signal. Phase modulations are visible only when sample contains nanoshells and when 808 nm laser is activated.
Although it clear from Fig. 5(d) that gold nanoshells can be detected by direct inspection of the measured phase under certain conditions, an enhanced SNR can be achieved by detecting modulations in the phase signal. One approach is to Fourier transform (FT) the measured phase and search for a peak at the precisely-known 808 nm modulation frequency. This concept is illustrated in Fig. 6. The plots in Fig. 6 show the frequency spectrum of the measured phase for each test condition in Fig. 5. In each case, the phase data from t=0 to t=4 ms was removed. The slow phase increase from gradual heating in Fig. 5(d) was removed prior to Fourier transformation by subtracting a quadratic fit from the measurement. When the 808 nm laser is disabled (Fig. 6(a,b)), whether the sample contained deionized water (Fig. 6(a)) or a nanoshells solution (Fig. 6(b)), the only characteristic feature of the frequency spectra is 1/f noise. The same is true when the 808 nm laser is modulated at 500 Hz but the sample contains deionized water (Fig. 6(c)). However, when the 808 nm laser is modulated at 500 Hz and the sample contains nanoshells, a strong peak is seen in the frequency spectrum of the measured phase at exactly 500 Hz (Fig. 6(d)). Smaller harmonic peaks are also visible at 1 kHz frequency increments, consistent with the FT of a triangular waveform repeating at 500 Hz.
Fig. 6. Fourier transform of phase vs. time curves, measured at back surface of cuvette. (a). Deionized water with 808 nm laser deactivated. (b). 1×1010 mL-1 nanoshell solution with 808 nm laser deactivated. (c). Deionized water with 808 nm laser modulated at 500 Hz. d, 1×1010 mL-1 nanoshell solution with 808 nm laser modulated at 500 Hz. Strong peak is observed at 500 Hz when nanoshells are present and the 808 nm laser is activated. SNR, signal-to-noise ratio.
The “signal” in this measurement is defined as the peak FT amplitude within ±20 Hz of the nominal 808 nm modulation frequency. This range was selected to allow for a small absolute error in the modulation frequency. The “noise” is defined as the peak FT amplitude within the same ±20 Hz window measured using the first 4 ms of data. During this time the 808 nm laser was disabled in all cases, allowing for an accurate and consistent estimate of phase noise at the 808 nm modulation frequency. The SNR measurements were repeated five times for each combination of sample type (with and without nanoshells) and 808 nm laser state (disabled or enabled for t>4 ms). The SNR is defined as the ratio of the signal peak to the noise value. In Fig. 6, SNR values are shown as the mean of five repeated measurements, plus or minus one standard deviation. As shown in Fig. 6(a,c), SNR values are insignificant when nanoshells are not present. This indicates the lack of a photothermal modulation signal. SNR values are also insignificant when nanoshells are present but the 808 nm laser is switched off (Fig. 6(b)). However, when nanoshells are present in the sample and the 808 nm laser is modulated at 500 Hz, the SNR is 131 ±91.
3.2 Effect of modulation frequency on signal to noise ratio
As the modulation frequency of the 808 nm laser is increased, the frequency of the phase modulation in the nanoshell solution also increases. This has the benefit of shifting the FT peak to higher frequencies where 1/f noise is reduced. However, the sample volume illuminated with the 808 nm beam has less time to heat and cool, resulting in a smaller optical path modulation and lower FT peak amplitude. To investigate this tradeoff, three sets of additional experiments were conducted using 808 nm modulation frequencies of 1, 15, and 60 kHz. Each experiment was repeated five times under the same conditions to test for consistency. The results are shown in Fig. 7. Figure 7(a–c) shows the measured phase from the nanoshell solution at one transverse position. The vertical line in each plot indicates the time when the 808 nm beam was switched on at t=4 ms. The insets show enlarged views of 500 µs (Fig. 7(b)) and 90 µs (Fig. 7(c)) segments of the measured phase, with clear modulations visible. For all three cases, a gradual path change is visible due to slow heating of the sample, although this effect is lower at 15 and 60 kHz.
Fig. 7. Measured phase (a–c) and Fourier transforms (d–f) of measured phase at various 808 nm laser modulation frequencies. Red lines in (a–c) show time when 808 nm laser was activated. Insets in (b,c) show enlarged views of measured phase. 808 nm laser modulation frequencies were 1 kHz (a,d), 15 kHz (b,e), and 60 kHz (c,f). SNR, signal-to-noise ratio.
Figure 7(d–f) shows the FT of the measured phases, beginning at 4 ms and with the slow phase component removed by subtracting a quadratic fit. The SNR is measured at each modulation frequency as described in 3.1 above. An optimum SNR of 112 ±45 is achieved at a modulation frequency of 15 kHz. At a modulation frequency of 1 kHz, the SNR suffers from low frequency 1/f noise near the baseband. At a modulation frequency of 60 kHz, the benefits of decreased 1/f noise are outweighed by the decrease in peak modulation amplitude. It is expected that the optimal modulation frequency will vary depending on the optical properties of the sample, nanoshell concentration, and 808 nm laser power level.
3.3 Effect of measurement time on signal to noise ratio
For in vivo imaging applications, it is important to determine how long one region of the sample must be measured in order to obtain a reasonable contrast agent SNR. Longer measurement times lead to increased SNR, but decreased overall frame rates. This tradeoff was evaluated by shortening the time window used in the FT for an 808 nm modulation frequency of 15 kHz. One representative dataset was used for this test, and the results are shown in Fig. 8. A linear relationship between the observation time and the SNR is observed in the measured data. As the observation time window is decreased from 28 ms to 2 ms, the SNR decreases linearly from 108 to 7. These results suggest that observation times of only a few ms may be required per transverse position in order to obtain reasonable nanoshell contrast using FT analysis. Other data analysis techniques may be developed in the future that can obtain similar contrast in shorter time periods.
Fig. 8. Measured gold nanoshell signal-to-noise ratio (SNR) as the sample observation time is decreased for an 808 nm laser modulation frequency of 15 kHz.
3.4 Thermal modeling
It is also important for future in vivo applications to determine the magnitude of the temperature change in the portion of the sample illuminated by the 808 nm laser. Since it is difficult to directly measure small temperature fluctuations occurring over milliseconds in a volume of ~3×10-3 mm3, the temperature change induced by the 808 nm laser was estimated using thermal modeling. The temperature change in the sample is estimated from the measured phase variations by modeling the effects of temperature on optical path length. Although the mechanical and thermal dynamics of the phantom used in these experiments are fairly complex, some understanding of the system behavior can be achieved by modeling two major effects that work to change the optical path length in opposing directions as the sample temperature is increased. First, the refractive index of water decreases with increasing temperature, which tends to decrease the optical path length. Second, the volumetric thermal expansion coefficient of water is positive near room temperature, which tends to increase the optical path length. For the phantom apparatus used in these experiments, the measured optical path length of the sample, z(T), varies with temperature T and can be modeled as:
Here, L(T) is the physical path length and n(T) is the refractive index. A change in optical path length Δz that occurs due to a change in temperature ΔT relative to an initial condition T 0 can be written as:
Here, z 0 is the initial optical path length at T 0. With a volumetric coefficient of expansion β, an initial physical path length L 0, an initial refractive index n 0, and a variation in refractive index with temperature dn/dT, the change in optical path length can be expressed using:
This formulation assumes that the fluid column illuminated by the 808 nm laser is free to expand only in the axial direction. Axial expansion could occur since the cuvette cover was not tightly fixed to the body. Note that the absolute change in optical path length associated with one thermal modulation is <120 nm for a 500 Hz modulation frequency and <7 nm for a 60 kHz modulation frequency. These small size scales complicate the dynamic response of the expanding and contracting water column, so the model used here may not precisely reflect the actual behavior of the system. In this model, dn/dT is assumed to be constant with temperature. The expression for Δz can be expanded to give:
The swept source OCT phase microscope measures phase changes Δϕ in the OCT interference fringes. Δϕ is related to Δz through the expression:
Therefore the estimated sample temperature variation can be calculated by solving the following quadratic expression for ΔT :
Although it is possible to explicitly solve Eq. (10) for ΔT, in reality β is a function of temperature as well, β=β(T). For water, the variation in β is significant and changes by ~50% between 20 and 30°C. Therefore Eq. (10) was solved numerically using a mathematics package (Matlab, The Mathworks, Inc.).
Figures 9(a), 9(b) show the estimated sample temperature T in the illuminated volume versus the observed OCT signal phase change Δϕ for a room temperature of 20°C. For this model, values of dn/dT=-91×10-6 °C-1, T 0=20 °C, L 0=200 µm, and λ 0=1315 nm were used. β values ranged from 207×10-6 °C-1 at 20°C to 385×10-6 °C-1 at 40°C. Since β is several times larger than dn/dT, sample expansion dominates the system and the net optical path length changes is positive. The quadratic term in Eq. (10) does not contribute substantially to the solution but was taken into account nonetheless. Figure 9(a) shows the T vs. Δϕ curve for a phase change of 0–10 rad, while Fig. 9(b) shows a range of 1–2 rad. Experimental results showed average phase modulations of ±575, ±369, ±93, and ±32 mrad at 808 nm laser modulation frequencies of 0.5, 1, 15, and 60 kHz, respectively, as shown in Fig. 5(d) and Fig. 7. By comparing the measured phase modulations to the model results shown in Fig. 9(a), the estimated temperature fluctuations are ±1.47, ±0.98, ±0.26, and ±0.09°C at 808 nm laser modulation frequencies of 0.5, 1, 15, and 60 kHz, respectively. If the 808 nm laser is held at one transverse position for 26 ms, slow phase increases of ~4–8 rad are observed, corresponding to temperature increases of ~7.4–12.3°C. This slow temperature increase can be minimized in in vivo imaging applications by translating the beam faster, since only a few ms of observation time per transverse location may be required. Additionally, the use of a more tightly focused 808 nm beam would permit more rapid cooling, as discussed below. This result also highlights the potential of this technique for conducting photothermal therapy as well as OCT imaging, since even larger temperature gradients could be induced by increasing the 808 nm exposure levels.
Fig. 9. Thermal modeling results showing estimated sample temperature calculated from observed phase changes. a, Phase changes of 0–10 rad. b, Phase changes of 0–2 rad.
For in vivo applications, it is desirable to reduce optical exposure levels and prevent cumulative heating over time. Therefore the 808 nm power level should be minimized while maintaining a reasonable nanoshell SNR. A second model was applied to understand photothermal contrast behavior as the 808 nm power and spot size are scaled. The system was modeled as an absorbing Gaussian cylinder surrounded by an infinite and homogenous medium. This simplified model does not take into account the effects of scattering on the 808 nm beam shape, but provides a reasonable estimate of the relationship between induced temperature change and 808 nm beam parameters. The model also assumes that the spot size is much smaller than the penetration depth due to absorption, which is true for the parameter space considered below.
The heat conduction equation in a cylindrical geometry is given by:
Here, t is time, z is the axial distance from the top of the cylinder, r is radial distance from the center of the cylinder, µa is the absorption coefficient, φ is the fluence rate of the laser, ρ is the density of the medium, c is the specific heat of the medium, and α is the thermal diffusivity of the medium. For small spot sizes relative to the absorption depth, the heat conduction equation is dominated by radial heat transfer [51]. Combined with a cylindrical geometry, this allows the heat conduction equation to be solved in closed form. For one 808 nm modulation period, the temperature variation can therefore be modeled as [51]:
Here, E is the irradiance, W is the 1/e2 beam radius, α is the thermal diffusivity, and tp is the exposure duration. E and α are calculated from fundamental material properties of the sample, given by:
Here, PPLS is the pulse power of the 808 nm laser and k is the thermal conductivity of the surrounding medium.
Equations (12) and (13) were used to model the expected temperature profile for an 808 nm modulation frequency of 500 Hz over one pulse period. The following parameter values were used for the model: ρ=1000 kg/m3, c=4186 J/kg K, tp=1 ms, µa=388 m-1, and k=0.6 W/m K. The value for µa was obtained from measurements performed by the nanoshell manufacturer. Thermal responses were modeled for 808 nm beam radii of 5–40 µm. The pulse power was chosen to give a constant temperature increase of 1 °C for each beam radius, corresponding to average powers of 2.6–18.2 mW at radii of 5–40 µm. The results of the model are shown in Fig. 10(a), with the legend indicating the beam radius and pulse power used to generate each curve. During the second half of each modulation cycle (t=1 ms to t=2 ms), the simulation indicates that temperatures decrease to 18%–62% of their peak values. This suggests that a slow temperature drift, similar to that shown in Fig. 5 and Fig. 7, may be difficult to avoid. Decreasing the beam radius, however, enables a given thermal increase to be achieved with less incident power, and also results in more rapid cooling. This would result in a larger phase modulation and lower tissue exposure, emphasizing the need to reduce the 808 nm beam diameter for future in vivo applications.
Figure 10(b) shows a comparison between the model results and the thermal response estimated from the actual phase measurements shown in Fig. 5(d) for t=4 ms to 6 ms. The measured phase was converted to an estimated temperature increase using Eq. (10). The model parameters were adjusted to reflect the actual experimental conditions. A beam radius of 70 µm was used with an average incident power of 138 mW. The model and estimation show good correlation, indicating that the thermal models described here are reasonable.
Fig. 10. (a). Thermal modeling results showing expected temperature increases for 808 nm beam radii of 4–40 µm. b, Comparison of thermal model (red curve) and thermal response estimated from phase measurement (blue curve) for a modulation frequency of 500 Hz, beam radius of 70 µm, and average power of 138 mW.
4. Discussion
The results shown here describe a photothermal modulation method for detecting gold nanoparticles using OCT. Changes in the optical path length of a sample containing gold nanoshells are thermally induced using a modulated 808 nm laser diode. These phase changes are measured over time using a phase-sensitive, swept source OCT microscope. Extremely high SNR values are obtained using this modulation technique.
Although other techniques used to detect gold nanoparticles with OCT have used different phantom systems, which makes direct comparison difficult, reported SNR’s have typically ranged from 1.5–5 [29, 32–34, 36]. There has been one previous report of SNR values of 79–631 using highly backscattering nanoshells in a non-scattering water sample, but the SNR decreased to 5 when a scattering tissue phantom that more closely approximated biological tissue was used [32]. The phantom used in the photothermal experiments described here does not include scattering, but accurately reflects the absorption properties of biological tissue [52] and achieves mean SNR’s of 2–131. The photothermal modulation method is expected to perform well in scattering tissue since the technique is less sensitive to background noise because phase changes are detected at a specific modulation frequency away from baseband. In addition, this technique uses absorption rather than scattering to generate contrast. The high SNR also suggests that lower concentrations of nanoparticles could be detectable using this method than other methods, which is relevant for in vivo applications where the agent is administered systemically and accumulates at lower levels in diseased tissue [34].
Although good performance was achieved in this phantom experiment, additional studies are required to demonstrate the utility of the technique in vivo. The phantom used here was selected to test the concept of photothermal detection with the most critical optical property of the sample, absorption, isolated from other effects. Although there are challenges for applying this method in vivo, OCT phase microscopy has previously been applied for studies of living cells and tissue by placing a thin coverslip in contact with the sample [40, 53–55]. Similar approaches may be used for the technique described here, with the top surface of a thin glass window or endoscope sheath in light contact with the tissue providing the reference field. Other groups have demonstrated that optically-induced thermal gradients can result in physical tissue displacements ex vivo that are detectable with phase-sensitive OCT [56]. Although these previous studies induced large temperature gradients and measured correspondingly large phase changes, they indicate that the photothermal modulation contrast mechanism described here should remain valid in tissue using smaller temperature gradients and measuring smaller phase modulations.
More fundamentally, recent reports have indicated that it is difficult to accurately measure small phase changes <0.1 rad in vivo using OCT [57]. This phase uncertainty is due to speckle noise and sample motion, and can be reduced through spatial averaging and other signal processing techniques. With photothermal detection of gold nanoshells, phase variations of ±0.098 rad can be induced at a modulation frequency of 15 kHz, and much larger variations are possible at lower modulation frequencies. This indicates that observation of the photothermal modulation signal should be possible in vivo. The necessity to spatially average the phase measurements over the axial or transverse dimensions in order to detect the photothermal modulation could decrease the effective spatial resolution or imaging speed of this technique. It would still be possible, however, to obtain conventional 3D-OCT images at high resolution concurrently with the contrast-enhanced images by applying standard OCT signal processing.
Another challenge for in vivo applications is ensuring that the thermal changes induced by the 808 nm laser are small enough to preclude inadvertent tissue damage. The thermal changes have two components: a small-amplitude component associated with the 808 nm modulation over tens of microseconds, and a slow rise in temperature associated with gradual heating over tens of milliseconds. For this discussion, it is assumed that in vivo temperature increases will be similar to those obtained in the phantom experiments. The results shown here indicate that for optimal SNR conditions, the temperature rise associated with the 808 nm modulation is ~0.52 °C. This is an extremely small thermal change that is not expected to cause tissue damage. The gradual temperature increase associated with slow sample heating at optimal SNR conditions is ~8 °C over 26 ms. For in vivo applications, however, the observation time at one transverse spot could be reduced to 5 ms while maintaining an SNR of ~20 which would cause an overall temperature increase of only ~3.5°C. Since the tissue would only be exposed to this elevated temperature for several milliseconds as the beam is scanned in the transverse dimension, no significant tissue damage is expected. By comparison, photothermal and photodynamic therapies typically require temperature increases of several tens of degrees for many tens of seconds to induce permanent tissue damage. Observation times may be further decreased by developing new analysis algorithms that more efficiently detect the photothermal phase modulation. The use of a single mode laser diode instead of the multimode 808 nm laser diode would allow thermal gradients to be induced over a more localized area. This would reduce the power required to create a given phase modulation and would simultaneously increase cooling rates, further increasing safety margins.
One final consideration for in vivo applications is the effect of scattering in biological tissue. Scattering may decrease the effectiveness of the photothermal modulation technique for large tissue depths, since 808 nm light penetration will be reduced. In the phantom experiments performed here, the phase measurement was performed at a fluid/glass interface at a single transverse position. In an in vivo application, the OCT and 808 nm beams would be continuously scanned across the sample in a transverse direction. Due to the spatial distribution of scattering centers in biological tissue, phase measurements would be obtained at each axial position in the sample in a manner similar to Doppler OCT techniques. The photothermal phase modulation may be detected by comparing the measured phase at each axial point to the measured phase in subsequent axial lines, or by comparison to a fixed phase reference such as a glass window or endoscope sheath. The combination of transverse beam scanning and axially distributed phase measurements will add speckle noise to the photothermal signal. Since nanoshell contrast is derived from a phase modulation at a precisely known frequency, however, the effects of random background noise such as speckle and sample motion are expected to be minimized. Increasing the 808 nm modulation frequency may also have important benefits in scattering systems, since this will allow more phase modulation periods to be captured over each speckle cell. Higher modulation frequencies could therefore increase OCT frame rates and take maximum advantage of the speed advantage offered by FDML lasers.
Although the initial results shown here suggest that the photothermal modulation technique may provide performance benefits in vivo, more study is required to verify this. Future experiments will focus on evaluating SNR performance in solid tissue phantoms that more closely approximate the scattering, thermal conductivity, and mechanical properties of biological tissue. Minimum detectable nanoshell concentrations will be studied in these phantoms and assessed for benefits compared to scattering- or absorption-based contrast. Finally, experiments in in vitro and in vivo biological tissue are needed to validate the photothermal contrast technique.
5. Conclusions
We have demonstrated a new type of active contrast agent detection technique for OCT imaging. Gold nanoshells are detected by inducing photothermal modulation in a sample with an 808 nm laser diode. Absorption of modulated 808 nm light by the nanoshells results in an optical path length change and corresponding phase modulation. This modulation is measured using a swept source OCT phase microscope and detected at the modulation frequency using Fourier transform methods. The temperature gradients associated with the photothermal modulation are small, indicating that the technique should be safe for in vivo use. High nanoshell SNR’s are achieved, indicating that the detection of low nanoshell concentrations in vivo may be possible. For future applications in scattering systems such as biological tissue, detection of a phase modulation at a known frequency should significantly reduce background noise from speckle and sample motion. In the future, the use of molecularly-targeted nanoshell contrast agents with photothermal modulation OCT imaging may enable highly sensitive and specific detection of diseases such as cancer.
Acknowledgments
This research was sponsored in part by the National Institutes of Health R01-CA75289-11 and R01-EY11289-20; the Air Force Office of Scientific Research FA9550-040-1-0011 and FA9550-010-0046; the National Science Foundation BES-0522845 and ECS-0501478. Mr. Adler acknowledges support from the Natural Sciences and Engineering Research Council of Canada. Dr. Fujimoto receives royalties from intellectual property owned by MIT and licensed to LightLabs Imaging and Carl Zeiss Meditec. Dr. Huber receives royalties from intellectual property owned by MIT and licensed to LightLabs Imaging.
References and Links
1. D. Huang, E. A. Swanson, C. P. Lin, J. S. Schuman, W. G. Stinson, W. Chang, M. R. Hee, T. Flotte, K. Gregory, C. A. Puliafito, and J. G. Fujimoto, “Optical Coherence Tomography,” Science 254, 1178–1181 (1991). [CrossRef] [PubMed]
2. S. Yazdanfar, M. D. Kulkarni, and J. A. Izatt, “High resolution imaging of in vivo cardiac dynamics using color Doppler optical coherence tomography,” Opt. Express 1, 424–431(1997). [CrossRef] [PubMed]
3. Z. Chen, T. E. Milner, D. Dave, and J. S. Nelson, “Optical Doppler tomographic imaging of fluid flow velocity in highly scattering media,” Opt. Lett. 22, 64–66 (1997). [CrossRef] [PubMed]
4. Z. Ding, Y. Zhao, H. Ren, J. S. Nelson, and Z. Chen, “Real-time phase-resolved optical coherence tomography and optical Doppler tomography,” Opt. Express 10, 236–245 (2002). [PubMed]
5. V. X. D. Yang, M. L. Gordon, B. Qi, J. Pekar, S. Lo, E. Seng-Yue, A. Mok, B. C. Wilson, and I. A. Vitkin, “High speed, wide velocity dynamic range Doppler optical coherence tomography (Part I): System design, signal processing, and performance,” Opt. Express 11, 794–809 (2003). [CrossRef] [PubMed]
6. V. X. D. Yang, M. L. Gordon, E. Seng-Yue, S. Lo, B. Qi, J. Pekar, A. Mok, B. C. Wilson, and I. A. Vitkin, “High speed, wide velocity dynamic range Doppler optical coherence tomography (Part II): imaging in vivo cardiac dynamics of Xenopus laevis,” Opt. Express 11, 1650–1658 (2003). [CrossRef] [PubMed]
7. V. X. D. Yang, M. L. Gordon, S. J. Tang, N. E. Marcon, G. Gardiner, B. Qi, S. Bisland, E. Seng-Yue, S. Lo, J. Pekar, B. C. Wilson, and I. A. Vitkin, “High speed, wide velocity dynamic range Doppler optical coherence tomography (Part III): in vivo endoscopic imaging of blood flow in the rat and human gastrointestinal tracts,” Opt. Express 11, 2416–2424 (2003). [CrossRef] [PubMed]
8. D. C. Adler, T. H. Ko, P. R. Herz, and J. G. Fujimoto, “Optical coherence tomography contrast enhancement using spectroscopic analysis with spectral autocorrelation,” Opt. Express 12, 5487–5501 (2004). [CrossRef] [PubMed]
9. C. Y. Xu, P. S. Carney, and S. A. Boppart, “Wavelength-dependent scattering in spectroscopic optical coherence tomography,” Opt. Express 13, 5450–5462 (2005). [CrossRef] [PubMed]
10. S. D. Dyer, T. Dennis, L. K. Street, S. M. Etzel, T. A. Germer, and A. Dienstfrey, “Spectroscopic phasedispersion optical coherence tomography measurements of scattering phantoms,” Opt. Express 14, 8138–8153 (2006). [CrossRef] [PubMed]
11. D. J. Faber, E. G. Mik, M. C. Aalders, and T. G. van Leeuwen, “Light absorption of (oxy-)hemoglobin assessed by spectroscopic optical coherence tomography,” Opt. Lett. 28, 1436–1438 (2003). [CrossRef] [PubMed]
12. D. J. Faber, E. G. Mik, M. C. G. Aalders, and T. G. van Leeuwen, “Toward assessment of blood oxygen saturation by spectroscopic optical coherence tomography,” Opt. Lett. 30, 1015–1017 (2005). [CrossRef] [PubMed]
13. U. Morgner, W. Drexler, F. X. Kartner, X. D. Li, C. Pitris, E. P. Ippen, and J. G. Fujimoto, “Spectroscopic optical coherence tomography,” Opt. Lett. 25, 111–113 (2000). [CrossRef]
14. R. Leitgeb, M. Wojtkowski, A. Kowalczyk, C. K. Hitzenberger, M. Sticker, and A. F. Fercher, “Spectral measurement of absorption by spectroscopic frequency-domain optical coherence tomography,” Opt. Lett. 25, 820–822 (2000). [CrossRef]
15. C. Y. Xu, C. Vinegoni, T. S. Ralston, W. Luo, W. Tan, and S. A. Boppart, “Spectroscopic spectral-domain optical coherence microscopy,” Opt. Lett. 31, 1079–1081 (2006). [CrossRef] [PubMed]
16. C. Vinegoni, J. S. Bredfeldt, D. L. Marks, and S. A. Boppart, “Nonlinear optical contrast enhancement for optical coherence tomography,” Opt. Express 12, 331–341 (2004). [CrossRef] [PubMed]
17. Y. Jiang, I. Tomov, Y. M. Wang, and Z. P. Chen, “Second-harmonic optical coherence tomography,” Opt. Lett. 29, 1090–1092 (2004). [CrossRef] [PubMed]
18. B. E. Applegate, C. Yang, A. M. Rollins, and J. A. Izatt, “Polarization-resolved second-harmonic-generation optical coherence tomography in collagen,” Opt. Lett. 29, 2252–2254 (2004). [CrossRef] [PubMed]
19. S. Yazdanfar, L. H. Laiho, and P. T. C. So, “Interferometric second harmonic generation microscopy,” Opt. Express 12, 2739–2745 (2004). [CrossRef] [PubMed]
20. M. V. Sarunic, B. E. Applegate, and J. A. Izatt, “Spectral domain second-harmonic optical coherence tomography,” Opt. Lett. 30, 2391–2393 (2005). [CrossRef] [PubMed]
21. J. P. Su, I. V. Tomov, Y. Jiang, and Z. P. Chen, “High-resolution frequency-domain second-harmonic optical coherence tomography,” Appl. Opt. 46, 1770–1775 (2007). [CrossRef] [PubMed]
22. K. D. Rao, M. A. Choma, S. Yazdanfar, A. M. Rollins, and J. A. Izatt, “Molecular contrast in optical coherence tomography by use of a pump-probe technique,” Opt. Lett. 28, 340–342 (2003). [CrossRef] [PubMed]
23. B. E. Applegate and J. A. Izatt, “Molecular imaging of endogenous and exogenous chromophores using ground state recovery pump-probe optical coherence tomography,” Opt. Express 14, 9142–9155 (2006). [CrossRef] [PubMed]
24. Z. Yaqoob, E. McDowell, J. G. Wu, X. Heng, J. Fingler, and C. H. Yang, “Molecular contrast optical coherence tomography: a pump-probe scheme using indocyanine green as a contrast agent,” J. Biomed. Opt. 11, 054017 (2006). [CrossRef] [PubMed]
25. T. M. Lee, A. L. Oldenburg, S. Sitafalwalla, D. L. Marks, W. Luo, F. J. J. Toublan, K. S. Suslick, and S. A. Boppart, “Engineered microsphere contrast agents for optical coherence tomography,” Opt. Lett. 28, 1546–1548 (2003). [CrossRef] [PubMed]
26. C. Y. Xu, J. Ye, D. L. Marks, and S. A. Boppart, “Near-infrared dyes as contrast-enhancing agents for spectroscopic optical coherence tomography,” Opt. Lett. 29, 1647–1649 (2004). [CrossRef] [PubMed]
27. A. L. Oldenburg, J. R. Gunther, and S. A. Boppart, “Imaging magnetically labeled cells with magnetomotive optical coherence tomography,” Opt. Lett. 30, 747–749 (2005). [CrossRef] [PubMed]
28. C. Loo, A. Lin, L. Hirsch, M. H. Lee, J. Barton, N. Halas, J. West, and R. Drezek, “Nanoshell-enabled photonics-based imaging and therapy of cancer,” Technol. Cancer Res. T. 3, 33–40 (2004).
29. H. Cang, T. Sun, Z. Y. Li, J. Y. Chen, B. J. Wiley, Y. N. Xia, and X. D. Li, “Gold nanocages as contrast agents for spectroscopic optical coherence tomography,” Opt. Lett. 30, 3048–3050 (2005). [CrossRef] [PubMed]
30. J. Chen, F. Saeki, B. J. Wiley, H. Cang, M. J. Cobb, Z.-Y. Li, L. Au, H. Zhang, M. B. Kimmey, X. Li, and Y. Xia, “Gold nanocages: bioconjugation and their potential use as optical imaging contrast agents,” Nano Lett. 5, 473–477 (2005). [CrossRef] [PubMed]
31. A. L. Oldenburg, F. J. J. Toublan, K. S. Suslick, A. Wei, and S. A. Boppart, “Magnetomotive contrast for in vivo optical coherence tomography,” Opt. Express 13, 6597–6614 (2005). [CrossRef] [PubMed]
32. A. Agrawal, S. Huang, A. W. H. Lin, M. H. Lee, J. K. Barton, R. A. Drezek, and T. J. Pfefer, “Quantitative evaluation of optical coherence tomography signal enhancement with gold nanoshells,” J. Biomed. Opt. 11, 041121 (2006). [CrossRef] [PubMed]
33. A. L. Oldenburg, M. N. Hansen, D. A. Zweifel, A. Wei, and S. A. Boppart, “Plasmon-resonant gold nanorods as low backscattering albedo contrast agents for optical coherence tomography,” Opt. Express 14, 6724–6738 (2006). [CrossRef] [PubMed]
34. A. M. Gobin, M. H. Lee, N. J. Halas, W. D. James, R. A. Drezek, and J. L. West, “Near-infrared resonant nanoshells for combined optical imaging and photothermal cancer therapy,” Nano Lett. 7, 1929–1934 (2007). [CrossRef] [PubMed]
35. J. Oh, M. D. Feldman, J. Kim, H. W. Kang, P. Sanghi, and T. E. Milner, “Magneto-motive detection of tissue-based macrophages by differential phase optical coherence tomography,” Laser. Surg. Med. 39, 266–272 (2007). [CrossRef]
36. T. S. Troutman, J. K. Barton, and M. Romanowski, “Optical coherence tomography with plasmon resonant nanorods of gold,” Opt. Lett. 32, 1438–1440 (2007). [CrossRef] [PubMed]
37. C. M. Pitsillides, E. K. Joe, X. B. Wei, R. R. Anderson, and C. P. Lin, “Selective cell targeting with light-absorbing microparticles and nanoparticles,” Biophys. J. 84, 4023–4032 (2003). [CrossRef] [PubMed]
38. K. Sokolov, M. Follen, J. Aaron, I. Pavlova, A. Malpica, R. Lotan, and R. Richards-Kortum, “Real-time vital optical imaging of precancer using anti-epidermal growth factor receptor antibodies conjugated to gold nanoparticles,” Cancer Res. 63, 1999–2004 (2003). [PubMed]
39. C. Loo, A. Lowery, N. Halas, J. West, and R. Drezek, “Immunotargeted nanoshells for integrated cancer imaging and therapy,” Nano Lett. 5, 709–711 (2005). [CrossRef] [PubMed]
40. M. A. Choma, A. K. Ellerbee, C. H. Yang, T. L. Creazzo, and J. A. Izatt, “Spectral-domain phase microscopy,” Opt. Lett. 30, 1162–1164 (2005). [CrossRef] [PubMed]
41. C. Joo, T. Akkin, B. Cense, B. H. Park, and J. E. de Boer, “Spectral-domain optical coherence phase microscopy for quantitative phase-contrast imaging,” Opt. Lett. 30, 2131–2133 (2005). [CrossRef] [PubMed]
42. B. J. Vakoc, S. H. Yun, J. F. de Boer, G. J. Tearney, and B. E. Bouma, “Phase-resolved optical frequency domain imaging,” Opt. Express 13, 5483–5493 (2005). [CrossRef] [PubMed]
43. D. C. Adler, R. Huber, and J. G. Fujimoto, “Phase-sensitive optical coherence tomography at up to 370,000 lines per second using buffered Fourier domain mode-locked lasers,” Opt. Lett. 32, 626–628 (2007). [CrossRef] [PubMed]
44. V. P. Zharov and D. O. Lapotko, “Photothermal imaging of nanoparticles and cells,” IEEE J. Sel. Top. Quant. 11, 733–751 (2005). [CrossRef]
45. A. V. Brusnichkin, D. A. Nedosekin, M. A. Proskurnin, and V. P. Zharov, “Photothermal lens detection of gold nanoparticles: Theory and experiments,” Appl. Spectrosc. 61, 1191–1201 (2007). [CrossRef] [PubMed]
46. R. Huber, M. Wojtkowski, and J. G. Fujimoto, “Fourier Domain Mode Locking (FDML): A new laser operating regime and applications for optical coherence tomography,” Opt. Express 14, 3225–3237 (2006). [CrossRef] [PubMed]
47. R. Huber, M. Wojtkowski, K. Taira, J. G. Fujimoto, and K. Hsu, “Amplified, frequency swept lasers for frequency domain reflectometry and OCT imaging: design and scaling principles,” Opt. Express 13, 3513–3528 (2005). [CrossRef] [PubMed]
48. R. Huber, D. C. Adler, and J. G. Fujimoto, “Buffered Fourier domain mode locking: unidirectional swept laser sources for optical coherence tomography imaging at 370,000 lines/s,” Opt. Lett. 31, 2975–2977 (2006). [CrossRef] [PubMed]
49. D. C. Adler, Y. Chen, R. Huber, J. Schmitt, J. Connolly, and J. G. Fujimoto, “Three-dimensional endomicroscopy using optical coherence tomography,” Nature Photon. 1, 709–716 (2007). [CrossRef]
50. R. Huber, D. C. Adler, V. J. Srinivasan, and J. G. Fujimoto, “Fourier domain mode locking at 1050 nm for ultra-high-speed optical coherence tomography of the human retina at 236,000 axial scans per second,” Opt. Lett. 32, 2049–2051 (2007). [CrossRef] [PubMed]
51. M. J. C. van Gemert, G. W. Lucassen, and A. J. Welch, “Time constants in thermal laser medicine .2. Distributions of time constants and thermal relaxation of tissue,” Phys. Med. Biol. 41, 1381–1399 (1996). [CrossRef] [PubMed]
52. T. L. Troy and S. N. Thennadil, “Optical properties of human skin in the near infrared wavelength range of 1000 to 2200 nm,” J. Biomed. Opt. 6, 167–176 (2001). [CrossRef] [PubMed]
53. M. A. Choma, A. K. Ellerbee, S. Yazdanfar, and J. A. Izatt, “Doppler flow imaging of cytoplasmic streaming using spectral domain phase microscopy,” J. Biomed. Opt. 11, 024014 (2006). [CrossRef] [PubMed]
54. A. K. Ellerbee, T. L. Creazzo, and J. A. Izatt, “Investigating nanoscale cellular dynamics with cross-sectional spectral domain phase microscopy,” Opt. Express 15, 8115–8124 (2007). [CrossRef] [PubMed]
55. E. J. McDowell, A. K. Ellerbee, M. A. Choma, B. E. Applegate, and J. A. Izatt, “Spectral domain phase microscopy for local measurements of cytoskeletal rheology in single cells,” J. Biomed. Opt. 12, 044008 (2007). [CrossRef] [PubMed]
56. B. J. Vakoc, G. J. Tearney, and B. E. Bouma, “Real-time microscopic visualization of tissue response to laser thermal therapy,” J. Biomed. Opt. 12, 020501 (2007). [CrossRef] [PubMed]
57. S. Makita, Y. Hong, M. Yamanari, T. Yatagai, and Y. Yasuno, “Optical coherence angiography,” Opt. Express 14, 7821–7840 (2006). [CrossRef] [PubMed]
