molecular imaging is widely used in pre-clinical studies of the diseases that cause the greatest burdens of morbidity and mortality in the developed world (e.g., cancer, cardiovascular disease, diabetes, etc.) [1
]. When performed at high resolution, in vivo
molecular imaging can provide insight into the mechanisms of disease progression and drug resistance on a cellular level, thereby unraveling heterogeneities in pathological expression and drug response that are critical for understanding and combatting these diseases [2
]. Microscopy, including confocal and multiphoton microscopy, has been the standard for high resolution molecular imaging in live cells and tissues. However, these microscopy techniques suffer from relatively shallow imaging depths. Magnetic resonance imaging (MRI) and positron emission tomography (PET) have been the standard for functional imaging deep within the body, but these methods lack cellular-level resolution.
Optical coherence tomography (OCT) fills a niche between high resolution microscopy and whole body imaging techniques with cellular-level resolution and penetration depths in tissue that exceed the imaging depths of microscopy. This three-dimensional, non-invasive imaging technique provides an especially attractive scale for monitoring mouse models of disease. However, contrast in standard OCT images is based largely on differences in scattering cross section, which can be minimal amongst certain molecular species. Thus, augmenting standard OCT images with sensitive and specific molecular contrast represents an area of significant interest.
Functional extensions of OCT, including magneto-motive (MMOCT) [3
], spectroscopic [5
], pump-probe [8
], and photothermal OCT (PT-OCT), have demonstrated molecular contrast. Specifically, PT-OCT has recently received much attention [10
] for a number of reasons. First, PT-OCT is able to identify and separate absorbing targets from the scattering background through active detection of photothermal heating [20
] (which is also independent of tissue mechanical properties, unlike MMOCT). Second, PT-OCT is highly sensitive to absorbing targets in the sample due to lock-in detection and low background. Finally, PT-OCT can exploit the rapid recent advancements in nanotechnology to develop efficient, molecularly-targeted contrast agents. For example, gold nanoparticles with near infrared (NIR) plasmon resonance peaks have been investigated for imaging and photothermal therapy [21
] and are particularly attractive contrast agents for PT-OCT.
PT-OCT leverages the photothermal heating phenomenon, where photon absorption by an imaging target of interest (e.g., an absorbing nanoparticle) leads to a temperature change in the environment surrounding the target [20
]. These local temperature changes cause thermoelastic expansion of the sample and shifts in the local index of refraction [25
]. The photothermal-induced shifts in the local index of refraction and geometric path length alter the local optical path length (OPL, the product of index of refraction and geometric path length). OPL changes due to photothermal heating can be directly imaged via the phase information in an OCT interferogram. The PT-OCT signal has been shown to increase with increasing pump beam power and absorber concentration, and the PT-OCT signal to noise ratio increases with the number of repeated photothermal cycles [10
PT-OCT has previously been characterized and demonstrated in vitro
and ex vivo
with targeted gold nanospheres [10
], non-targeted gold nanoshells [11
], gold nanorods [14
], gold nanorose [13
], and carbon nanotubes [18
]. However, to date, no contrast agents have been imaged with PT-OCT in vivo
. Two studies have demonstrated in vivo
PT-OCT of hemoglobin (for quantifying blood oxygenation) [15
], although only point scans (not images) were collected over multiple second long acquisition times. The primary goal of this study is to demonstrate the ability of PT-OCT to image contrast agents in vivo
Gold nanorods (GNRs) are especially appealing contrast agents for in vivo
PT-OCT because of their resonance in the NIR tissue optical window (~650-900 nm), tunable optical absorption properties (based on their physical dimensions) [26
], and efficient absorption (compared to gold nanoshells and nanospheres) [27
]. GNRs are also on a more advantageous size scale for in vivo
molecular imaging (tens of nanometers in size) compared to nanoshells (hundreds of nanometers in size). Finally, the full width half max (FWHM) of the GNR absorption peak is much smaller than that of nanoshells [28
], which is desirable to avoid attenuation of the imaging beam by the contrast agent.
The idealized pairing of PT-OCT with GNR contrast agents could allow for three-dimensional in vivo molecular imaging in a currently unexploited regime of resolution and penetration depth. However, prior to in vivo molecular imaging, the PT-OCT signal must be characterized, optimized, and tested in simpler in vivo systems without the added complexity of targeted molecular imaging. In this paper, GNRs are demonstrated as robust PT-OCT contrast agents, achieving pM-scale concentration sensitivity. The PT-OCT signal is also characterized with respect to imaging speed, photothermal beam power, and OCT magnitude signal. In addition, experimental PT-OCT signals are directly compared to photothermal heating models. Finally, we demonstrate PT-OCT imaging of GNRs in phantoms as well as the first documented in vivo images of contrast agents using PT-OCT.