MolecularOpticalCoherenceTomography Contrast ...user.physics.unc.edu/.../Oldenburg-2015-MCOCT-chapter.pdf1 contrast agents in MRI and echogenic microbubbles in ultrasound. In fact,

Molecular Optical Coherence TomographyContrast Enhancement and Imaging 47Amy L. Oldenburg, Brian E. Applegate, Jason M. Tucker-Schwartz,Melissa C. Skala, Jongsik Kim, and Stephen A. Boppart

47.1 Introduction

Histochemistry began as early as the nineteenth century, with the development of

synthetic dyes that provided spatially mapped chemical contrast in tissue [1]. Stains

such as hematoxylin and eosin, which contrast cellular nuclei and cytoplasm,

greatly aid in the interpretation of microscopy images. An analogous development

is currently taking place in biomedical imaging, whereby techniques adapted for

MRI, CT, and PET now provide in vivo molecular imaging over the entire human

body, aiding in both fundamental research discovery and in clinical diagnosis and

treatment monitoring. Because OCT offers a unique spatial scale that is intermedi-

ate between microscopy and whole-body biomedical imaging, molecular contrast

A.L. Oldenburg (*)

Department of Physics and Astronomy and the Biomedical Research Imaging Center, University

of North Carolina at Chapel Hill, Chapel Hill, NC, USA

e-mail: [email protected]

B.E. Applegate

Department of Biomedical Engineering, Texas A&M University, College Station, TX, USA

J.M. Tucker-Schwartz • M.C. Skala

Department of Biomedical Engineering, Vanderbilt University, Nashville, TN, USA

J. Kim

Department of Electrical and Computer Engineering, Bioengineering, Medicine, and the Beckman

Institute for Advanced Science and Technology, University of Illinois at Urbana-Champaign,

Champaign, IL, USA

S.A. Boppart

Biophotonics Imaging Laboratory, Beckman Institute for Advanced Science and Technology,

University of Illinois at Urbana-Champaign, Urbana, IL, USA

Departments of Bioengineering, Electrical and Computer Engineering, and Medicine, University

of Illinois at Urbana-Champaign, Urbana, IL, USA

# Springer International Publishing Switzerland 2015

W. Drexler, J.G. Fujimoto (eds.), Optical Coherence Tomography,DOI 10.1007/978-3-319-06419-2_48

1429

mailto:[email protected]

OCT (MCOCT) also has great potential for providing new insight into in vivo

molecular processes. The strength of MCOCT lies in its ability to isolate signals

from a molecule or contrast agent from the tissue scattering background over large

scan areas at depths greater than traditional microscopy techniques while

maintaining high resolution.

MCOCT involves the use of OCT image acquisition and/or processing tech-

niques to generate image contrast using endogenous molecular species or exoge-

nous molecular probes of interest [2]. It should be noted that several other chapters

in this book cover techniques that provide information about tissue composition

that will not be repeated here, including nonlinear interferometric vibrational

imaging (NIVI), second harmonic generation OCT, and optical coherence

elastography. This chapter does address spectroscopic OCT (SOCT) in the context

of chromophore and probe detection, while the reader should refer to related SOCT

chapters in this book for greater detail on SOCT processing methods.

This chapter is structured according to the physical mechanisms used for contrast,

starting from direct schemes whereby probes add to or subtract from the optical

backscattering spectrum that comprises the OCT signal (i.e., scattering- and

absorption-based contrast, respectively), then covering methods that indirectly mod-

ulate the OCT signal (i.e., pump-probe, magnetomotive, and photothermal OCT). It

should be noted that any scheme developed for MCOCT must be compatible with

interferometric detection, which precludes the use of several physical mechanisms

such as fluorescence emission. In this chapter, imaging technology development will

be emphasized and presented with selected examples of biomedical applications.

Interestingly, many endogenous and exogenous probes can be sensed by more than

one method. For example, photothermal contrast relies upon absorption, and there-

fore, the same agents provide contrast in both photothermal and spectroscopic OC-

T. The development of exogenous imaging probes (or contrast agents) that enable

MCOCT is also a rich and varied topic that poses particular challenges in materials

science and targeted delivery. Here we introduce a variety of molecular probes in the

context of specific MCOCT imaging strategies, while the reader is referred elsewhere

[3–6] for detailed information about probe development.

47.2 Scattering-Based Contrast

The native “contrast” observed in OCT is light that has been coherently

backscattered. All MCOCT methods are based upon modifying this backscattering

signal in a way that can provide additional, molecular information about the

sample. One of the most straightforward strategies is to increase the OCT signal

directly with a probe particle that exhibits high backscattering. Analogous methods

include the use of positive T1 contrast agents in MRI and echogenic microbubbles

in ultrasound. In fact, microbubbles also happen to have reasonably high

light scattering and were one of the first types of contrast agents studied with

OCT [7]. Oil-filled protein microspheres were subsequently found to offer flexibil-

ity in loading the shell or core with nanoparticles to further increase the optical

1430 A.L. Oldenburg et al.

scattering [8], and variants of these are in continued use with magnetomotive OCT,

as discussed below [9]. It is important to note, however, that any new contrast agent

must undergo thorough safety and efficacy testing before it can be used on humans,

such as that required by the Food and Drug Administration in the United States.

As such, the study of agents already approved for human use can more readily lead

to clinical translation. Interestingly, it was recently found that several commonly

used ophthalmic medications provide scattering-based OCT contrast [10]. As

shown in Fig. 47.1, OCT reveals the diffusion of several medications within the

anterior chamber after administration. They were also shown to enhance the

visibility of corneal incisions postoperatively, which may provide a method for

assessing wound integrity. Future adaptations of molecularly targeted agents may

further broaden the functionality of OCT in ophthalmology.

While scattering-based contrast agents are readily visible within the highly trans-

parent anterior segment of the eye, the ability to detect these types of agents endo-

scopically or on the skin is more challenging, as they must be distinguishable against

the already high optical scattering of the tissue.Mie theory provides exact solutions for

light scattering from spherical particles, shelled spheres, and spheroids [11].

As a general rule, there is a rapid increase in scattering with particle size (scattering

cross section, ss / d6) in the Rayleigh regime (d << l) and a rapid increase in

Lipid-based Artificial Tears

Control

Triamcinolone

Prednisolone

Fig. 47.1 OCT images of

fresh, ex vivo porcine eyes

after intra-cameral injection

of contrast agents. From topto bottom: Control,triamcinolone acetonide

40 mg/mL, prednisolone

acetate 1 %, and lipid-based

artificial tears. The agents

provide positive contrast

within the anterior chamber

that has a different pattern

depending on the type of

agent used. These differences

are expected because the

agents are comprised of

varying particulate or oil-in-

water suspensions, with

differing mean particle sizes,

reflectivities, and diffusivities

47 Molecular Optical Coherence Tomography Contrast Enhancement and Imaging 1431

scattering as the material refractive index is different (higher or lower) than that of the

(typically aqueous) medium. At the same time, one must weigh the choice of the

material and the particle size against the biocompatibility and the ability for the probes

to access their target, respectively, for the needed application.

One of the most useful types of OCT contrast agents, which will be discussed

many times throughout this chapter, is plasmonic gold nanoparticles in their various

forms (nanospheres, nanoshells, nanorods, etc.). This is because gold is highly

unreactive and consequently relatively biocompatible, while at the same time

providing a surface plasmon resonance (SPR) effect at the red and near-infrared

wavelengths used in OCT [12]. This SPR is evident as either a strong absorption or

scattering spectral peak, with a transition from predominantly absorbing (low

albedo) to predominantly scattering behavior (high albedo) as the particle size is

increased; for spheres, the transition occurs at d � 80 nm [12]. Nanoshells in

particular have been highly developed for scattering-based OCT contrast [13–15].

They are comprised of a silica core and gold shell and offer spectral tunability by

adjusting the core diameter and shell thickness [3]. Figure 47.2 displays a demon-

stration of enhanced OCT contrast in the tumors of mice systemically intravenously

Tumor Tissue + PBSc d

a b

Min

200 µm

Max

Normal Tissue + PBS

Tumor Tissue + Nanoshells

Normal Tissue + Nanoshells

Glass

Skin

Muscle

Glass

Skin

Tumor

Fig. 47.2 OCT images of tissues from mice with subcutaneous tumors systemically treated with

phosphate-buffered saline (PBS) as a control (a, c) and multifunctional nanoshells (b, d).Enhanced retention of nanoshells in the tumor in panel (d) provides better delineation of tumor

borders, as well as subsequent tumor-specific photothermal ablation (Reprinted with permission

from Ref. [13]. Copyright 2007 American Chemical Society)


injected with nanoshells [13]. These results highlight the enhanced permeation and

retention (EPR) effect, whereby the permeable vasculature of tumors acts to trap

particles, providing selective targeting [16]. Importantly, EPR further enables

site-targeted treatment; in this example, the nanoshells were designed to be

nearly equally light scattering and absorbing, providing both imaging contrast

(via scattering) and photothermal therapy (via absorption).

It should also be noted that, in cases where the SPR is narrow compared to the

bandwidth of the light used in OCT, it may be possible to employ spectroscopic OCT

(SOCT) techniques to distinguish the SPR signature, providing enhanced specificity

against the tissue scattering background. While this idea has been explored [17],

a confounding factor that makes this method challenging is the highly modulated

backscattering spectrum typically obtained from Mie scatterers (d � l) within the

tissue. In current practice, SOCT techniques are much more commonly employed to

detect absorption-based contrast agents, which is the focus of the following section.

47.3 Absorption-Based Contrast

Light absorption is a very attractive molecular process to exploit for contrast, both

because of the potential signal strength and because essentially all molecular species

have the capacity to absorb light. The imaging light used in OCT is spectrally broad,

and hence, the backscattered spectrum may be utilized to identify the absorption

spectrum of endogenous or exogenous species present within the tissue. The group of

techniques designed to extract this information are collectively referred to as spec-

troscopic optical coherence tomography (SOCT) [18].

The different algorithms developed for SOCT diverge in how they deal with the

trade-off between spatial and spectral resolution. One approach is to use multiple

light sources to collect independent OCT images with different center wavelengths.

Relatively straightforward algorithms such as spectral triangulation [19] may then

be implemented to extract the depth-resolved backscattered spectrum. A judicious

choice of center wavelengths can facilitate the detection of highly peaked spectral

features with limited spectral resolution. This approach has largely been used to

detect dyes such as indocyanine green (ICG) [19].

An alternate approach is to directly utilize the broad spectral bandwidth of the

light source and use the short-time Fourier transform (STFT) to gain spectral

resolution at the expense of spatial resolution [20]. This approach has the advantage

that it is entirely a post-processing technique; hence, it can be tailored to maximize

contrast to a target contrast agent. Likewise, the time-frequency distribution (TFD)

need not be limited to the STFT but may optimized as well to maximize the spatial

and spectral resolution [21]. Wax and coworkers [22] have recently developed an

algorithm that incorporates two STFTs, one with a narrow spectral window and one

with a broad spectral window. The two TFDs are multiplied point by point to

generate a TFD with both high spatial and spectral resolution.

Exogenous chromophores for SOCT are largely repurposed, commercially

available fluorescent dyes. Utilizing these dyes carries with it the advantage of


awealth of biological and chemical research aimed toward targeting particular disease

states of tissue, chemical species, or morphologies. Some examples include ICG [23],

photodynamic therapy-related dyes [24], and fluorescent microspheres [25].

Dyes typically also have strongly peaked spectra which enable detection via fairly

simple methodologies. For instance, a commercial NIR absorbing dye (H.W. Sands,

ADS7460) which exhibits a sharp absorption peak at 740 nm was used to produce

contrast in an 800 nm OCT system by effectively clipping the shorter wavelengths,

resulting in a redshift of scattered light [26].

The major endogenous chromophore is hemoglobin. Detection of hemoglobin

absorption with SOCT has been explored as a method to measure blood oxygen

saturation [27]. Wax and coworkers [22] recently measured the oxygen saturation

along with fluorescein dye injected into the bloodstream in a mouse window

chamber model using METRiCS OCT which uses the two TFD methods noted

above along with OCT imaging at nontraditional wavelengths. Their imaging

bandwidth spans the 455-695 nm range which overlaps strong peaks in the oxy-

and deoxyhemoglobin spectrum. Selected results from this work demonstrating the

endogenous and exogenous tissue contrast as well as SO2 measurements are shown

in Figs. 47.3 and 47.4.

Plasmonic gold nanoparticles have also been widely employed for absorption-

based contrast, where the SPR peaks are tuned within the OCT imaging band by

varying the size of specific geometrical features of the nanoparticles. For instance,

light-absorbing gold nanorods are tuned by varying their aspect ratio (length over

width) while maintaining a length typically <100 nm to favor absorption over

scattering. Many of the different particle geometries have been explored for

contrast in SOCT, including gold nanospheres [28], gold nanorods [28, 29], and

gold nanocages [30]. For example, gold nanorods were imaged after injection into

excised human breast carcinoma tissue [29]. As mentioned above, light-absorbing

Fig. 47.3 Conventional OCT (a) and METRiCS OCT (b) images, located above point (e) in the

en face (x–y) image in Fig. 47.4.White x and z scale bars, 100 mm (Reprinted with permission from

Ref. [22]. Copyright 2011 Macmillan)


gold nanoparticles are, at the same time highly effective for photothermal

cancer therapy, where a high power laser is used to irradiate particle-laden

tumors [31]. The synergy between imaging and therapy, which allows us to monitor

permeation and diffusion of SPR particles into tissues before treatment, aids in

particle development for improved delivery and informs the design of more effec-

tive treatment protocols.

47.4 Pump-Probe OCT

Pump-probe optical coherence tomography (PPOCT) is fundamentally the fusion

of optical coherence tomography with pump-probe absorption spectroscopy.

Spatially resolving the pump-probe interaction can provide molecular contrast

for absorbing agents in a tissue sample similar to SOCT. The major advantages of

PPOCT are that contributions to the OCT light attenuation due to scattering and

absorption are easily separated and there is no compromise between spatial and

spectral resolution. The disadvantages are that the optical setup is more complicated

and typically requires pulsed light sources. Only under special circumstances could

PPOCT be accomplished with a swept OCT laser source or superluminescent diode.

Fig. 47.4 (a) En face METRiCS OCT image with arrows indicating points where the spectra areextracted. White x and y scale bars, 100 mm. (b–e) Spectral profiles from corresponding points in

(a). Measured spectral profiles (black) are superposed with the theoretical oxy- (dashed red) anddeoxy- (dashed blue) hemoglobin normalized extinction coefficients, and normalized absorption

of NaFS (dashed green). Also shown are the SO2 levels and the relative absorption of NaFS with

respect to total hemoglobin (e � NaFS/Hb). All spectra were selected from depths immediately

below each corresponding vessel (Reprinted with permission from Ref. [22]. Copyright 2011

Macmillan)


A typical PPOCT system has the following features: The probe is the light in the

sample arm of the OCT interferometer, i.e., the same light used for OCT imaging

serves as the probe light. A separate pump beam co-propagates with the OCT light in

the sample arm. The pump is typically amplitude modulated at frequency f0. Transferof the modulation onto the backscattered probe (OCT) signal at f0 is then evidence ofabsorption of the probe light by some tissue absorber. In time-domain OCT

implementations, the PPOCT signal appears as sidebands on the Doppler carrier

frequency, fD � f0 [32]. In spectral-domain OCT, the PPOCT signal can be extracted

from an M-scan by Fourier transformation along the time axis (at each depth) and

filtering around f0. For the process to work, absorption of the pump by the contrast

agent must change the absorption/scattering properties at the probe wavelength.

The first experimental realization of PPOCT [33] demonstrated imaging of

methylene blue, a dye used for chromoendoscopy [34]. The specific physical

mechanism leading to PPOCT signal from methylene blue is well understood and

therefore serves as a germane example, where the energy level diagram is illus-

trated in Fig. 47.5a. The pump light drives a transition from the singlet ground

state (S0) to the first excited electronic state (S1). Molecular population in the

excited singlet state is transferred to the triplet state via a particularly efficient [35]

spontaneous process (S1 ! T1, t1-1). Methylene blue in its triplet state has

a resonant transition peaked at 830 nm (T1 ! T2). When the pump is on, an

830 nm probe can be absorbed by methylene blue, but when the pump is off, there

is no probe absorption. In reality the excited triplet state has a finite lifetime (analogous

to fluorescence lifetime) that is a function of the oxygen level in its local environment,

but varies from �200 ns to over 1 ms. Consequently, the pump and probe need not be

incident on the sample at the same time, but may be delayed in time by some fraction

of the excited state lifetime. Measurement of this characteristic lifetime may be used to

help differentiate among multiple chromophores. An example decay for methylene

blue is in Fig. 47.5b. The average decay time (lifetime) from T1 to S0 (t0�1) was

calculated to be 247 ns via tavg ¼ ∑ S � t/∑S, where S is the PPOCT signal at delay

time t [36]. In addition to the lifetime, the absorption spectrum at the pump or probe

may be measured by recording the PPOCT signal as a function of the pump or probe

wavelength, respectively.

Several molecular species in addition to methylene blue have been imaged using

PPOCT. Phytochrome A, a naturally occurring molecular switch which may be

reversibly optically pumped from one isomeric state to another, was imaged

in a tissue phantom [37]. Hemoglobin was measured in the gill filament arteries

of a zebrafish (Brachyrerio danio) using a time-domain 532 nm PPOCT system

with a 532 nm pump [32]. The same system was also used to image the fluorescent

protein DsRed in a transgenic zebrafish. Melanin was imaged in the first spectral-

domain PPOCT system in a phantom made from human hair embedded in chicken

breast tissue [38]. Melanin has also been imaged using a time-domain optical

coherence microscopy system [39]. Recent work [36] has demonstrated volumetric

imaging of microvasculature in Xenopus laevis using a two-color (532 nm pump,

830 nm probe) PPOCT system. Representative PPOCT cross sections overlain

on the OCT cross sections are shown in Fig. 47.6 along with volumetric


reconstructions of the vasculature measured with PPOCT. They have also demon-

strated the use of the characteristic lifetime to differentiate PPOCT signals from

two different chromophores.

Figure 47.7 shows a pair of capillary tubes loaded with methylene blue/micro-

spheres and bovine whole blood in heparin. The top panel (a) is the standard OCT

image showing similar signal from both capillary tubes. A PPOCT image with 2 ns

pump-probe delay (Fig. 47.7b) appears very similar to the OCT image. However,

when the pump-probe delay is increased to 24.8 ms, the signal from the methylene

blue/microsphere-loaded capillary tube decays, leaving only signals from the blood-

filled capillary. Taking advantage of the difference in lifetime between two chro-

mophores is an effective strategy for imaging multiple chromophores with PPOCT.

Future research in PPOCTmay lead in several directions. There is a clear potential

for imaging vasculature. While Doppler-based OCT can also measure vasculature,

a major advantage of PPOCT is that the signal is independent of the angle of flow,

while the Doppler signal approaches zero when the illumination is orthogonal to

Fig. 47.5 (a) Molecular energy level diagram for the methylene blue PPOCT mechanism. Driven

transitions are indicated by straight arrows and spontaneous transitions as zigzag arrows.(b) Measured normalized decay of the PPOCT signal due to methylene blue as a function of

delay between the pump and probe pulse. The decay has a characteristic average lifetime of 247 ns

(Modified and reprinted with permission from Ref. [36]. Copyright 2013)


the flow. The PPOCT signal is also molecularly specific, so it may be possible to

differentiate oxy- and deoxyhemoglobin and develop a PPOCT-based measure of

blood oxygen saturation. Furthermore, the imaging of exogenous contrast agents such

as methylene blue could potentially be used to tag and image specific molecular

species that are otherwise invisible to OCT. Such applications hinge on the demon-

stration of sufficient sensitivity eitherwithmethylene blue or some other discovered or

engineered contrast agent.

Fig. 47.6 (a–c) PPOCTB-scans overlaid on the

corresponding co-registered

OCT B-scans. Xenopus laevisvasculature is clearly

depicted. Arrows in (a) pointto capillaries that were not

visible in conventional OCT.

(d) Volumetric

reconstructions of the

microvasculature (Modified

and reprinted with permission

from Ref. [36]. Copyright

2013)


47.5 Magnetomotive OCT

47.5.1 Theory and Instrumentation

Magnetomotive OCT (MMOCT) is a method for contrasting the distribution of

magnetic particles based on their induced motion within a temporally modulated,

magnetic field gradient [40]. Figure 47.8 illustrates the mechanism of MMOCT,

showing howmagnetic particles inside the imaging volume are mechanically pulled

toward an electromagnet placed in the imaging arm of an OCT system. Typically,

phase-sensitive OCT is then used to track the motion of light scattering tissue

structures that are mechanically coupled to the particles [41]. Modulation of

the electromagnet thus leads to phase modulation that can be band-pass filtered

at the modulation frequency to detect the magnetomotion. Because human

tissues are only very weakly magnetic (magnetic susceptibility |w| < 10�5),

Fig. 47.7 (a) OCT B-scan of

methylene blue (MB) andhemoglobin (Hb) in acrylic

capillary tubes. (b)Corresponding PPOCT

B-scans with a pump-probe

delay of 3 ns. (d)Corresponding PPOCT

B-scan with a pump-probe

delay of 24.8 ms. The absenceof the MB tube demonstrates

the potential for using the

lifetime as an effective

method for differentiating

multiple chromophores


MMOCT provides high specificity against the tissue background, on the order

of 105 when using probes of w � 1 [42].

A class of biomedical imaging probes currently used in MRI, called superpar-

amagnetic iron oxides (SPIOs), are ideal for MMOCT because they are designed to

exhibit large w, avoid irreversible aggregation that is associated with ferromagnetic

agents, and are composed of iron oxide which has a proven safety profile.

FDA-approved MR liver contrast agents such as Feridex™, for example, have

been shown to provide excellent MMOCT contrast [43]. Another type of

MMOCT probe is protein microspheres encapsulating SPIO-containing ferrofluid,

which then offer flexibility in adding targeting ligands and therapeutic

payloads [44].

Implementing MMOCT on an existing phase-sensitive OCT system is relatively

straightforward. A small electromagnet can be placed on either the same side

(as shown in Fig. 47.8) or opposite side of the tissue to provide a magnetic field

gradient oriented along the imaging axis. Somewhat counterintuitively, the strength

of the magnetic field should only be on the order of 0.1T; higher fields will typically

saturate the magnetic particles and reduce the detection sensitivity [41, 43].

The absolute sensitivity of MMOCT can be determined by considering the balance

of forces between the diamagnetic tissue, which is pushed away from the magnet,

and from the paramagnetic particles, which are pulled toward the magnet.

Fig. 47.8 Mechanism of magnetomotive contrast in OCT. A solenoid placed in the imaging arm

of an OCT system provides a magnetic gradient force, F!, on magnetic particles inside tissue

according to the gradient of the magnetic field, ∇ B!, and the magnetization and volume of the

particles, M!and V, respectively. The resultant elastic displacement of mechanically coupled light

scattering structures, Dz, is sensed as a phase shift in the OCT interferogram, Df. w is the particle

magnetic susceptibility, m0 is the vacuum permeability, and n is the tissue refractive index at the

imaging beam wavelength, l


For a typical SPIO particle, the minimum particle concentration needed to tip this

force balance in favor of motion toward the magnet is on the order of 10 mg Fe/g.

Another important consideration is the elastic property of the tissue medium.

Magnetic particles in liquid do not undergo a restoring force during magnetic

field modulation, moving only in one direction, and exhibit little contrast by

conventional band-pass-filtered MMOCT. In a solid medium, the compliance of

the tissue dictates the amount of displacement Dz, resulting in MMOCT contrast

that is weighted by both the local particle concentration and the local tissue

stiffness. Owing to the nanoscale displacement sensitivity afforded by phase-

sensitive OCT systems, the tissue stiffness is typically of little detriment to the

overall MMOCT sensitivity, and sensitivities as low as 27 mg Fe/g have been

reported in optomechanical tissue phantoms [41]. The high sensitivity and speci-

ficity afforded by MMOCT have recently led to several new molecular imaging

application areas, which will be reviewed below.

47.5.2 MMOCT of Atherosclerosis

Atherosclerosis is a disease in which an arterial vessel wall thickens as a result

of the accumulation of fatty materials, including macrophages. Atherosclerosis is

promoted by low-density lipoproteins (LDL) and cholesterol (crystals) and results

in calcification in advanced lesions [45]. Standard intravascular OCT imaging

has been extensively investigated for applications in cardiology such as imaging

intraluminal 3D structure and function at high resolution, evaluating arterial

stents, and visualizing atherosclerotic plaque [46–51]. The addition of contrast

agents for use with intravascular OCT can enable site-specific molecular cardio-

vascular imaging, just as has been shown for ultrasound imaging using gas-filled

microbubbles [52]. Targeted contrast agents may enable the early detection

and localization of atherosclerotic lesions which may not be clearly evident in

structural OCT imaging or in other imaging modalities. Therefore, the combina-

tion of intravascular OCT and targeted molecular contrast enhancement with

MMOCT can potentially improve the sensitivity of early atherosclerotic lesion

detection. Figure 47.9 shows representative MMOCT images from an ex vivo

hyperlipidemic rabbit aorta. The RGD (arginine-glycine-aspartic acid)-

functionalized protein microspheres [44] were fabricated to target the aVb3integrin overexpressed in atherosclerotic lesions [53]. These microspheres

were loaded with SPIOs and a fluorescent dye, enabling multimodal imaging

using MMOCT, MRI, ultrasound, and fluorescence imaging. The ex vivo aorta

sample was perfused in a custom-designed flow chamber at physiologically

relevant pulsatile flow rates and pressures. The functionalized microspheres

have been successfully targeted to fatty streaks present during early-stage

atherosclerosis. The future development of a MMOCT catheter or a new

solenoid configuration for current commercial intravascular OCT systems

may enable in vivo MMOCT imaging of atherosclerosis-targeted magnetic

probes.


47.5.3 MMOCT of Breast Cancer

Breast cancer is one of the most commonly occurring cancers in women and has

a widespread effect on our society [54, 55]. Like other diseases, early detection is

the key in the treatment of breast cancer. Breast cancer screening methods include

manual clinical and self-breast exams, mammography, genetic screening,

ultrasound, and MRI. Breast cancer cells often overexpress aVb3 integrin and

HER-2/neu receptors which are considered to be biomarkers for targeted cancer

treatment [56]. Figure 47.10 shows an important step in breast cancer diagnosis

using MMOCT to contrast HER-2/neu-targeted SPIOs [57]. In this study, the

authors used a nitroso-methyl-urea (MNU) carcinogen-induced rat mammary

tumor model. In vivo MMOCT images of tumors from rats injected with targeted

SPIOs, nontargeted SPIOs, and saline exhibit an accumulation of SPIOs only in the

tumors of rats injected with targeted MNPs.

Fig. 47.9 Representative MMOCT and corresponding fluorescence confocal microscopy images

of hyperlipidemic rabbit aortas after administration of RGD microspheres. Parametric MMOCT

images are displayed showing the magnetomotive signal in green and the OCT signal in red. TheMMOCT signal in the targeted microsphere group was statistically significantly higher (p < 0.01)

than the nontargeted and control groups. Yellow lines in the aorta photos correspond to the imaging

locations. The dotted blue and red boxes are magnified to show the presence of individual

microspheres (white arrows). Scale bars are consistent across each row


Fig. 47.10 In vivo (a) MMOCT and (b) OCT of rat mammary tumors. The magnetomotive signal

(green) is superposed on the OCT (red) in MMOCT images. Prussian blue (PB) sections of (c, d)tumors and (e, f) livers from rats after injection with (left) targeted SPIOs, (center) nontargeted

SPIOs, and (right) saline. PB sections in (d, f) at 40� from boxed regions in (c, e), at 10�.

(g) Immunohistochemical-stained sections of (left) tumor from a targeted SPIO injected rat,

(center) tail injection site from a targeted SPIO injected rat, and (right) tumor from

a saline injected rat (Reprinted with permission from Ref. [57]. Copyright 2010 National Academy

of Sciences)


Furthermore, aVb3 integrin is also overexpressed in cancer cells [58]. As in the

previous section on atherosclerosis, the protein shell of microspheres can be

functionalized with the RGD tripeptide to target the aVb3 integrin. A recent study

reported MMOCT imaging of a rat mammary tumor containing RGD-targeted

magnetic protein microspheres [9]. MMOCT imaging was performed ex vivo on

the tumor approximately 4 h post injection of RGD-functionalized protein micro-

spheres. It was found that the aVb3 integrin-targeted microspheres preferentially

accumulated in the tumor.

These preliminary results demonstrate the feasibility of using targeted magnetic

nanoparticles and microspheres to detect breast cancer in OCT images. In addition,

the liquid-core protein microspheres can potentially also be used to deliver a drug

payload, such as the anticancer drug Taxol [59].

47.5.4 Platelets as Functional MMOCT Contrast Agents forThrombosis

Another way of achieving molecular specificity is to label cells with imaging agents

in vitro and to subsequently monitor the activity of the cells after in vivo admin-

istration. Platelets are cellular fragments present in blood that are responsible for

primary hemostasis (plug formation) as part of the blood clotting process [60].

Platelets respond to factors expressed when the blood vessel endothelium is dam-

aged, such as during trauma and atherosclerosis [61]. Therefore, platelets labeled

with SPIOs can be considered novel contrast agents for targeting localized endo-

thelial damage.

Interestingly, platelets have a unique method for nanoparticle uptake,

dubbed “covercytosis” [62], which is thought to be part of the body’s native

immune mechanism. As such, they avidly take up particles at levels reaching

hundreds [63], if not thousands [64], of femtograms of iron per platelet.

Platelets harvested from blood can also be partially fixed and lyophilized to

allow for long-term storage; these are known as rehydratable, lyophilized (RL)

platelets [65].

To date, in vitro and ex vivo studies have demonstrated the potential for

MMOCT to provide targeted imaging of reactive vascular sites using SPIO-labeled

RL platelets [43, 63]. In an ex vivo study, pig arteries were cannulated and

flowed with whole blood containing both native platelets and SPIO-RL platelets

at equal concentrations [43]. The endothelium of one set of arteries was lightly

injured, resulting in blood clotting at the injury sites. As shown in Fig. 47.11,

subsequent MMOCT of the luminal wall of these arteries revealed specific

magnetomotive contrast to injured artery only, due to the incorporation of

SPIO-RL platelets into the clots. These findings may be broadly translatable for

assessing internal bleeding and a broad spectrum of cardiovascular diseases such as

atherosclerosis.


47.6 Photothermal OCT

Photothermal OCT (PTOCT) provides sensitivity and specificity of OCT to

absorbers in a sample through active detection of photothermal heating. Photon

absorption by an endogenous chromophore or exogenous contrast agent leads to

a temperature rise in the surrounding environment. These local temperature

changes cause thermoelastic expansion of the sample and shifts in the local index

of refraction [66, 67], which in turn cause changes in the local optical path length.

These small, typically nanometer-scale, local optical path length changes can be

resolved with phase-sensitive OCT. Unlike fluorescence-based imaging modalities,

most absorption-based contrast agents do not undergo photobleaching, allowing for

constant PTOCT signal over time, even at high irradiance.

47.6.1 Theory and Instrumentation

In PTOCT, a separate laser source for photothermal heating is incorporated into the

sample arm of the OCT system, either via direct, free beam coupling at the sample

arm optics or through shared sample arm fiber optics. The heating laser wavelength

is chosen to match the peak absorption of the desired imaging target. In most

applications of PTOCT, square or sine wave amplitude modulation of the heating

beam (e.g., using a mechanical optical chopper or acousto-optical modulator) is

performed during temporal sampling (M-mode scanning) of each spot. An example

PTOCT instrumentation diagram can be seen in Fig. 47.12, with the incorporation

of an amplitude-modulated 808 nm laser into the sample arm of a spectral-domain

Fig. 47.11 Representative MMOCT images of ex vivo porcine arteries after exposure to SPIO-

labeled RL platelets in a flow chamber, revealing specific contrast to injured vascular endothelium.

Arteries were subsequently longitudinally cut and are imaged with the luminal wall facing upward.

Inset: TEM image of an SPIO-labeled platelet containing hundreds of SPIOs in its surface-

connected open canalicular system


Fig.47.12

(a)Experim

entalsetupofthePTOCTsystem

,wherePCdenotesthepolarizationcontroller.(b)Diagram

ofthedataprocessingmethodusedto

imagesentinel

lymphnodes

withPTOCT(A

daptedwithpermissionfrom

Ref.[68].Copyright2011American

Chem

ical

Society)


OCT system [68]. Amplitude modulation of the heating beam allows for digital

lock-in techniques to be used during signal processing, which can detect and isolate

the active heating dynamics from the passive scattering signal. Modulation fre-

quencies as low as 25 Hz [69] and as high as 120 kHz [70] have been reported in

PTOCT and PTOCM applications.

In PTOCT, the signal is isolated from an oversampled M-mode scan by

obtaining the Fourier transform (in the time dimension) of the OCT phase data at

each point in depth. The PTOCT signal is then defined as the magnitude of this

Fourier-transformed phase data at the modulation frequency. More complex signal

processing considerations are often taken into account to remove artifacts, includ-

ing fifth- [69] or sixth- [71] order polynomial background subtraction of the phase

data to minimize 1/f noise, baseline subtraction of nearby frequency components in

the FFT data to account for the additive noise floor in the signal [69, 71–73], and

averaging of overlapping short-time Fourier transforms over the M-mode scan to

better estimate the noise floor [73]. Previous investigations into the PTOCT imag-

ing parameters have demonstrated that the PTOCT signal increases linearly

with both absorber concentration [67–69, 72–75] and photothermal laser power

[69, 73, 74], decreases logarithmically with increased amplitude modulation fre-

quency [73], and has a constant mean value but increased noise level in the presence

of weak reflections in the sample [73].

47.6.2 PTOCT Applications

PTOCT has been demonstrated in vitro, ex vivo, and in vivo both with

endogenous and exogenous forms of contrast. The most common use of PTOCT is

to visualize ultralow concentrations of highly absorptive contrast agents in tissue and

in vitro. In particular, PTOCT can exploit the rapidly advancing field of nanotech-

nology to image contrast agents with strong, wavelength-tunable absorption peaks.

Gold nanoparticles are the most common PTOCT contrast agents due to their

biocompatibility and well-established surface modification chemistry, which can

allow for specific imaging of molecular targets. In vitro molecular imaging of

60 nm diameter epidermal growth factor receptor (EGFR)-targeted gold nanospheres

has been performed in agarose phantoms using a 532 nm heating laser at 25 Hz

modulation, where EGFR+ cells exhibited a 300 % increase in PTOCT signal

compared to EGFR cells [69]. This study also demonstrated a sensitivity of 14 ppm

(w/w) to gold nanospheres within a scattering background. Although this is the only

demonstration of true molecularly targeted imaging using PTOCT to date, a number

of applications have used nontargeted NIR-resonant nanoparticles in systems ranging

from phantoms to in vivo. Early investigations into PTOCT were performed with

gold nanoshells having an SPR tuned to the near-infrared [67]. In subsequent studies

with ex vivo human breast tissues, direct injections of 120 nm silica core and 16 nm

gold shell nanoshells at a 5 � 109/mL concentration were visible 300–600 mm deep

into the tissue using PTOCT with an 808 nm laser modulated at frequencies from 5 to

20 kHz, with 22mWof power on the sample [71]. Gold nanorods have also generated


interest as PTOCT contrast agents due to their tunability (based on the aspect ratio)

and particularly narrow SPR peak. Gold nanorods coated in poly(ethylene glycol)

(PEG) were found to have a significantly enhanced PTOCT signal at as low as 1 pM

concentration using 50 Hz modulation of an 808 nm laser interfaced with a 1,310 nm

OCT system [68]. The same system was used to image nonspecific uptake of gold

nanorods in sentinel lymph nodes (SLN). After injection with PEG-coated gold

nanorods, SLNs were dissected at varying time points from sacrificed mice and

imaged ex vivo after being embedded in 1 % agar gel. PTOCT was able to identify

the accumulation of nanorods within several SLN structures (Fig. 47.13, [68]).

Fig. 47.13 (a) Three-dimensional OCT projection image of a dissected sentinel lymph node

(SLN) at 48 h after gold nanorod injection. (b) 3D OCT view of SLN morphology with a cross-

sectional cut at a depth of 240 mm below the top surface. (c) 3D PTOCT view of SLN

corresponding to the cross-sectional cut displayed in (b) reveals structures within the SLN. (d)Schematic diagram and photograph of a dissected SLN. (Volume size ¼ 2.5 � 2.5 � 2.0 mm

(x–y–z)) (Adapted with permission from Ref. [68]. Copyright 2011 American Chemical Society)


A separate study was able to image, in vivo, picomolar concentrations of PEG-coated

gold nanorods directly injected into a mouse ear [73].

Applications of PTOCT are not limited to gold nanoparticles, but include

any targets that have strong absorption at the wavelength of the heating laser.

PTOCT has also been demonstrated using carbon nanotubes [72], indocyanine

green encapsulated poly(lactic-co-glycolic) acid nanoparticles [76], gold-iron

oxide nanoroses [75], and iron oxide-silica-gold multifunctional nanoprobes [77].

PTOCT has also been used to characterize hemoglobin oxygen saturation using

dual-wavelength approaches to probe the absorption differences of oxy- and

deoxyhemoglobin in vitro and in vivo [78, 79].

47.6.3 Recent PTOCT Advances

Recent advances in the field of PTOCT have been working toward providing

quantitative molecular imaging. First, measurements of the local slope in the

axial dimension of a PTOCT image has found some success in correcting for

phase accumulation in PTOCT [74]. Second, a method to perform photothermal

optical lock-in optical coherence microscopy (poli-OCM) has been developed for

real-time photothermal imaging [70]. In poli-OCM, one creates phase modulations

in the reference arm that are matched to the amplitude modulation frequency of the

heating beam, resulting in a demodulated photothermal signal of the absorbers in

the sample. The signal due to scatterers is then isolated from the photothermal

Fig. 47.14 Demonstration of the contrast selective to gold nanoparticles offered by poli-OCM, as

compared to dfOCM. A square lattice of isolated 40 nm gold particles on a glass surface immersed in

intravenous perfusion fluid, imaged with dfOCM (a), and poli-OCM (b). (d, e) correspond to cross

sections along the lines indicated in (a, b). Graph (c) depicts the signal along the lines in (a, b), while(f) corresponds to the axial signal along the line highlighted in (d, e). Scale bars: 10 mm (Adapted

with permission from Ref. [70]. Copyright 2012 Optical Society of America)


signal by setting the CCD integration time to a multiple of the modulation period.

This provides real-time photothermal imaging without the need for temporal

(M-mode) sampling or extensive digital processing. Pache et al. demonstrated

single particle detection of gold nanoparticles using poli-OCM with modulation

frequencies of 120 kHz while rejecting the scattering signal captured from their

dark-field optical coherence microscopy (dfOCM) system (Fig. 47.14, [70]).

Photothermal optical lock-in has yet to be demonstrated with a traditional OCT

system, but the underlying principles remain the same.

PTOCT is a promising imaging technique for isolating absorbers in a scattering

sample and thus provides specific and sensitive molecular imaging of both endog-

enous and exogenous contrasts. With recent advances in PTOCT optimization,

demonstrations in ex vivo and in vivo samples, and incorporation of optical lock-

in techniques, PTOCT promises to be not only sensitive and specific, but also a fast

method for imaging absorptive contrast agents in tissue.

47.7 Conclusion

It should be evident to the reader that MCOCT can be accomplished by

a wide variety of methods that each offer unique advantages, all of which

should be weighed when considering a specific biomedical application. All of the

methods presented here are in continued development, and we can expect to see

continued progress (greater sensitivity, specificity, imaging speeds, and higher reso-

lution) in the years to come. Also, new methods continue to emerge, such as imaging

thermal diffusion of micro- and nanoprobes [80, 81], which may allow one to sense

the local macromolecule concentration in bodily fluids such as mucus and blood.

Fundamentally, OCT provides a unique and flexible platform for exploring new

concepts in molecular imaging. Tying these imaging technology advances with

concomitant advances in nanotechnology and targeted delivery makes this a rapidly

changing field with many new opportunities for scientific discovery.

Acknowledgments Wewish to thank our many colleagues and collaborators conducting research

in this area and apologize that, due to length restrictions, we were unable to highlight more results.

We acknowledge Justis Ehlers at the Cole Eye Institute, Cleveland Clinic, for aiding in Fig. 47.1.

Some of the studies reported in this chapter were supported in part by a grant from the US National

Institutes of Health (R01 EB009073, S. A. B.).

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