Photodynamic therapy (PDT) initiates a photochemical reaction, which causes tumor destruction either by direct tumor cell killing or indirectly via vascular damage and an elicited immune response. For the photochemical reaction to occur, photosensitizer, oxygen, and light must be present in the target area. After administration, the kinetics of the distribution of the photosensitizer are tissue type- and site-dependent. The mechanism of action of PDT depends on both the photosensitizer and on the drug light (D-L) interval.
Although the idea of combined vascular and tissue damage is becoming more widely accepted, most clinical protocols for PDT are currently based on the assumption that optimum intervals between photosensitizer administration and illumination relate to the maximum difference between drug uptake in tumor and surrounding normal tissue. However, if vascular mediated tissue damage is the main determinant of tumor destruction, the optimal time interval should be more closely related to high blood plasma drug levels. Although several studies report on the effects of PDT on tissue and vasculature, the relationship between extent of vascular damage and efficacy of treatment is still not fully understood. 1, 2, 3, 4, 5, 6 Noninvasive methods that allow studying temporal changes in tissue perfusion following PDT could help to clarify the importance of vascular damage in clinical PDT. They could also serve as an early predictor of the efficacy of treatment.
Optical coherence tomography (OCT) is a high-resolution, cross-sectional imaging to technique that enables noninvasive examination of superficial tissue structures. By combining amplitude and phase-sensitive detection, it is possible to extract the spectral content of the backscattered light. When the light is reflected by a moving particle, the spectrum will acquire a frequency (Doppler) shift. OCT therefore enables cross-sectional imaging of blood flow in superficial vessels, in addition to structural imaging. The clinical use of OCT for imaging of morphological changes following PDT has been described for ophthalmological applications and for monitoring laser surgery of laryngeal carcinoma. 7, 8, 9, 10
In this study, the possibilities and limitations of optical coherence tomography for the assessment of PDT effects will be investigated on both normal and tumor tissue. For validation of the technique, several conditions will be imposed to influence the perfusion in a controlled way. Then, OCT is used to noninvasively obtain real-time information on the development of phototoxic tissue damage following mTHPC-PDT in normal and tumor tissue of mice.
OCT Measurement and Signal Processing
Measurements were performed using a fiber-based time-domain OCT setup with a 1300-nm super luminescent diode (SLD). Light from the SLD was split into two arms by a 50:50 fiber splitter. Longitudinal scanning was obtained by using a tilting mirror in the rapid scanning optical delay (RSOD) line in the reference arm.11 Transversal scanning was obtained by scanning the sample with the object beam. Signals at the detector were digitized and processed by a computer to form 2-D tomographic images ( , , axial resolution: , lateral resolution ).
For the results presented here, 32 consecutive axial samples were used to calculate a frequency spectrum using a Fourier transform algorithm (see Fig. 1 ). A Gaussian waveform was fitted to the spectrum to obtain a center frequency, amplitude, and spectrum width. The Fourier window was moved over the whole dataset (2-D area). A shift of the center frequency relates to flow or other causes of movement of the sampled volume. To correct for low-frequency motion artifacts, caused by breathing and beating of the heart, the average Doppler shift in a specific depth scan (a-scan) was calculated and subtracted from all values in that a-scan. Noise in the resulting dataset was further reduced by applying a moving average window ( pixels) and a threshold before the Doppler shift image was superimposed on the amplitude picture. The value of the applied threshold, obtained with control measurements described ahead, determined the minimal detectable flow.
Because of the averaging procedures and the unknown angle of the blood vessels with respect to the probing beam, no attempt was made to calculate real flow velocities. The direction of the flow was not taken into account. From the resulting images, the relative perfused area (RPA), which we define as the ratio of the number of pixels indicating perfusion over the total number of pixels in a region of interest, was determined. The values reported here are ratios of the RPA before and after the treatment and therefore represent relative changes in perfusion (RCP).
Where possible, three areas were identified in the images: skin, connective tissue, and tumor periphery. Figure 2 shows an OCT image in relation to a typical histology image of untreated mouse skin and subcutaneous tumor. A grayscale is used to visualize the differences in back scattering; a brighter layer indicates higher back scattering. The upper layers of skin (stratum corneum, epidermis, and dermis) are highly reflecting (area A in Fig. 2). Between the skin and the tumor or muscle, a low reflecting layer can be observed, which consists of loose connective tissue containing blood vessels (area B). Area C is a combination of a thin, highly vascularized muscle layer that is usually distinguishable below the connective tissue layer forming the capsule of the tumor and the upper part of the tumor.
To compare the PDT effects quantitatively, the slope of the backscattered light as a function of depth in the tissue was determined. The attenuation coefficient changes if the viability or configuration of the cells changes.2 As described by van Leeuwen, this slope is a function of both the point spread function (PSF) of the OCT system and the optical properties (attenuation coefficient) of the tissue.13 We have not determined the PSF of our system but fixed the focus at the air tissue boundary, thereby keeping the influence of the PSF constant throughout the measurements. Furthermore, a low NA lens was used to focus the beam on the tissue, which reduces the influence of the PSF on the signals. The reported values are therefore proportional to the attenuation coefficient and can consequently be used to follow relative temporal changes in tissue morphology.
Female nude BALB/c mice were used for all experiments. HNXOE cells (kindly provided by G. van Dongen, Vrije Universiteit, Amsterdam, The Netherlands) originated from a human head and neck squamous cell carcinoma metastasis in the oral cavity. HNXOE cells were cultured in Dulbecco modified Eagle medium (Gibco, Rockville, MD) supplemented with 5% fetal calf serum (Gibco) and antibiotics and grown at in a 5% incubator. To propagate donor tumors in vivo, HNXOE cells, in a volume of PBS, were injected subcutaneously in the lower dorsum of mice. When these donor tumors had reached a mean diameter of , tumor fragments were transplanted into the lower dorsum of recipient experimental mice. This two-step procedure was used to create well-localized spherical subcutaneous tumors with predictable growth velocity. The time required for the tumor to get to 5–7-mm diameter was about 27 days. During the imaging procedures, the mice were anesthetized by i.p. administration of a mixture of hypnorm and dormicum ( fluanisol fentanyl citrate and midazolam hydrochloride). To maintain body temperature, a temperature-controlled underlay was used.
Four control groups were used, each comprising two animals. In the first control group, the blood flow in the skin, connective tissue, and tumor was completely blocked by fixing a metal clamp around the tumor area (see Fig. 3 ). The effects on flow before, during, and after these interventions were measured by OCT. In groups 2 and 3, vasodilation was induced by administration of nicotinic acid ( solution) or hydralazine ( solution), respectively. Both drugs are vasoactive but have different working mechanisms. A vasodilation in the cutaneous blood vessels, resulting in an increased blood flow, was expected for both drugs. Nicotinic acid was also expected to lead to an improvement in tumor perfusion by preventing temporary shutdown of tumor vessels. 14, 15, 16, 17 Hydralazine is known to reduce the peripheral vascular resistance by directly relaxing arteriolar smooth muscle. This causes a relative decrease in perfusion of some tumor types, as the blood flow is preferentially diverted to critical normal tissues; the so-called steal effect.18, 19 In the fourth group, two animals received the photosensitizer but were not illuminated to assess the effect of the drug alone. In one animal of this last group, we imaged a needle that was inserted in the tumor and under the overlying skin. The OCT image was compared with the corresponding histology slide to identify the different layers in our measurement volume (data not shown).
mTHPC (Quanta Nova LTD, Scotland, supplied as a solution, drug in a mixture of 1:1:2 propylene glycol:ethanol:water) was administered via the tail vein at a dose of . At either or after administration, the mice were anesthetized, after which the skin of the flank and lower dorsum overlying the tumor was observed by OCT. After these pre-treatment measurements, the animals were illuminated with light. A total energy of was delivered using a bare fiber. The illumination spot was adjusted to cover the tumor with 0.5-cm margins. Repeated OCT measurements were made for up to 4 hours after illumination. The measurements were done in triplicate (for each time point, on each animal, three independent measurements) and three mice were used in each group. After the last measurement, the animals were sacrificed and tissue samples were collected from both treated and nontreated areas and processed for histology.
OCT measurements made before and 10 minutes after clamping of the tumor and overlying skin (group 1) clearly showed a dramatic drop in RPA after clamping (Fig. 4 ). Flow immediately-stalled in all three layers. After removal of the clamp, perfusion remained absent in both skin and tumor during the observation period (10 minutes). In the connective tissue layer, however, Doppler shifts above threshold were seen after removal of the clamp.
Perfusion of peripheral areas of the HNXOE xenograft, overlying connective tissue, and skin was also measured repeatedly after administration of nicotinic acid or hydralazine (Fig. 5 ). Both drugs induced an increase in perfusion in the skin immediately after administration, but changes induced in skin were much more pronounced for nicotinic acid (Fig. 5a). As expected, an increase in the tumor perfusion was observed after administration of nicotinic acid and a decrease after administration of hydralazine (Fig. 5b). Baseline values of perfusion were restored at 50 minutes and 120 minutes after administration of nicotinic acid and hydralazine, respectively. The same figures (Figs. 5a and 5b) show the average values of the RPS measured in skin of control group 4 (photosensitizer, no light). No significant changes were observed, which is also true for the RPS in the other layers.
Figure 6 shows the OCT images and corresponding histology before (Fig. 6a) and after PDT (Fig. 6b) using a 3 hr D-L interval. The two graphs on the right side show the intensity of the backscattered light as a function of depth.
The most striking histological difference in tissue structure after PDT was the large amount of edema that accumulated in the connective tissue layer (compare Figs. 6a and 6b). This caused an increase in scattering (see also Fig. 7 ) and a thickening of the connective tissue layer (Fig. 6b), which hampered imaging of the underlying tumor.
Changes in light attenuation after PDT for the 3 hr D-L (A) and 48h D-L (B) interval groups are shown in Fig. 7. All layers showed a trend for increased attenuation after PDT. This was greatest in the peripheral tumor layer for the 3hr D-L interval and greatest in the skin for the D-L interval, but none of the individual measurements was significantly different from baseline. The large error bars reflect biological variability, inhomogeneities in the layer, and the high sensitivity of the fitting procedure for the length of the segment over which the fit is done.
Changes in RCP with time after PDT are shown in Fig. 8 . After PDT with a short D-L interval perfusion of skin and tumor periphery decreased to values of 50% (skin) and 30% (tumor) after 160 minutes. Perfusion in the connective tissue initially increased, followed by a decrease in perfusion to a value that was approximately half the baseline value at 160 to 200 minutes. After PDT using longer D-L interval , no significant changes in RCP were seen.
The present study shows the possibilities and limitation of optical coherence tomography for the assessment of PDT effects on both normal and tumor tissue. For validation of the technique, several conditions were imposed to influence the perfusion in a controlled way. First the flow was completely stopped in the target area by clamping the tumor and skin. The rapid and complete inhibition of perfusion was used to optimize the algorithms used to extract the flow parameters from the raw data. Next, nicotinic acid or hydralazine were administered to check that tissue layer-dependent changes in perfusion could be followed with OCT.
The OCT tissue images that were taken after PDT showed a clear change in morphology (edema formation) and perfusion. These results were more pronounced for a D-L interval compared with a D-L interval. Furthermore, OCT could identify the different layers of tumor and overlying skin, with a maximal penetration of in this model. Flow information could be obtained over the full depth range of the structural image. The limited depth penetration represents one of the major limitations of OCT imaging. Motion artefacts in the signal, caused by breathing, heart beat, and muscle twitching were in the order of the local Doppler signal. Because this Doppler shift was usually present in the whole depth scan it could be taken into account in the algorithms and separated from a localized Doppler frequency.
Like ultrasound, OCT provides images that require interpretive skills and training. It is hard to draw conclusions from the images alone. Most toxic processes, e.g., apoptosis and necrosis, cause changes in optical attenuation. By quantifying the signals these processes can be followed in time. The trends in attenuation that we observed suggest treatment-related processes. We have previously shown that in cell suspensions apoptosis caused an increase in scattering while necrosis caused a decrease in scattering.12 However, in an in vivo situation additional processes will influence the attenuation of light, like the influx of edema and a change in blood volume. Further research on this subject is necessary.
Noninvasive imaging of pathologic tissue is preferable to invasive biopsies in many clinical conditions, not only for establishing a diagnosis but also for real-time assessment of treatment-related effects and as guidance during surgery, e.g., for ensuring tumor-free excision edges.
Several factors determine the clinical value of an imaging technique. The clinical value of OCT for the application described in this paper is determined by the imaging resolution, imaging depth, and possibility to assess functional data from the tissue. Considering the first two, for most imaging techniques a trade-off has to be made between imaging depth and resolution. In that respect, OCT can be placed between (fluorescence) confocal microscopy (CM) and high-resolution ultrasound (US). 20, 21, 22, 23 CM is capable of visualizing cell layers at micrometer resolution at the cost of a limited penetration depth of around . The state-of-the-art ultrasound currently has resolutions down to 30 micrometer resolution and a penetration depth up to .
In addition to CM and US, a number of new technologies are now available for noninvasive optical diagnostic imaging in conjunction with PDT, e.g., MRI, which has a spatial resolution of and is intrinsically sensitive to movement (flow).24 And, if assessment of perfusion is the main goal, laser Doppler perfusion imaging is an option.25, 26
The above-mentioned noninvasive techniques are only able to measure average perfusion over a large volume or do not have a high enough resolution to provide any detailed information on PDT effects in specific tissue layers.27 OCT might overcome these problems due to its high-resolution, depth-resolved measurements. A lot of research has been performed to evaluate the use of OCT to determine functional parameters, like flow, both in-vitro and for various clinical applications. 7, 28, 29, 30, 31, 32, 33, 34, 35, 36, 37 The downside of the technique is obviously the limited penetration of the light. This restricts the clinical use of OCT to superficial tissue layers. In this respect, the mouse model that was used in our experiments turned out to be very challenging. Because of the many strong reflecting layers, it was difficult to extract information from deeper layers in the tissue. Despite this limitation, we were able to follow the phototoxic effect on morphology and perfusion in this particular animal model application, and we may carefully conclude that Fig. 8 suggests that a 3-hr interval may be preferred over a 48-hr interval since there are more pronounced perfusion effects. Our findings justify further exploration of the technique. The concept of using noninvasive OCT measurements for real-time monitoring of morphology and perfusion may have applications in monitoring and optimization of anti-angiogenic or antivascular (cancer) therapy.
This study was supported by the Dutch Cancer Society, Grant NKI 2000-2282, and Stichting Dioraphte.