The challenge of characterizing the seemingly random geometric relationships between the visible biological features of tissues can be simplified by examining the Fourier spatial frequency (FSF) spectrum of the image.1 Drug delivery to tumors is well known to be chaotic and limited, partly from the perfusion limitations of dysfunctional neovasculature, but also because of the microscopic regional variations in composition which occur.2,3 However, while the phenomenological features of drug delivery limitations are known, understanding ways to mitigate the problem remains an ongoing research challenge. The characterization of this seemingly random microheterogeneity is an issue that would benefit significantly from automated methods to quantify delivery efficacy. With this as a goal, in this study, the features of tumor sections were analyzed for anisotropic structures and how this relates to flow kinetics, by examining the FSF spectrum of the image features. This letter presents the conceptual development of how structure could be used to automatically classify transport in systems of well-characterized tracers and tumor types.
The intensity map of any image can be decomposed into the superposition of a series of sinusoidal frequencies, with units of cycles per unit length, similarly to the way electrical signals can be decomposed into individual temporal frequencies. Recently, Russell et al.4 used FSF analysis to identify the characteristic spatial frequencies of dispersions of self-assembled biological macromolecules in different states of aggregation. Pu et al.1 demonstrated the use of FSF to distinguish between histological images of normal tissue and early-stage cervical cancer. Reufer et al.5 in a related application developed a method to incorporate directionality into FSF analysis, for identification of anisotropy in nanoparticle dynamics. Heterogeneous distribution of exogenous agents to solid tumor is thought to be associated with local tumor stromal heterogeneity: the relationship between nonuniform porosity from stromal growth and channel formation and nonuniform distribution of therapeutic or diagnostic agents has been observed both by histology,6 and in vivo MRI.7 The ultrastructure of the tumor stroma has been identified as a regulator of this phenomenon and is well known to be a major barrier in certain tumors, such as pancreatic cancer.8 Regions of invasive cell growth, disordered collagen, or high glocosaminoglycan content also can create a restricted geometry that defines a tortuous flow environment.9
In this study, we take advantage of the fact that in fluorescent tracer experiments, a region of poor tracer distribution demonstrates evidence of a barrier to transport. Using intensity maps of fluorescent tracers, we identify regions of marked fluorescence intensity anisotropy and then apply a directional averaging of the FSF over the intensity map of a conventionally stained histological section of that region. When a standard hematoxylin and eosin (H&E) stain is used, the eosin component, which stains protein, forms a pattern in the interstices between cells, i.e., an image of the protein distribution in the stroma. The characteristic frequency spectrum recovered from the intensity of the eosin staining is a quantitative record of how finely or coarsely spaced the structural features revealed in the image are. In this letter, we show that the characteristic spatial frequency spectrum of intratumor protein associated with directional anisotropy of tracer distribution is a unique identifier of transport barriers and establishs a systematic method of quantifying the relationship between tumor structure and drug distribution.
Methods and Materials
Images were from cut sections of xenograft tumors grown in nude mice. All animal procedures were approved by the Dartmouth Institutional Animal Care and Use Committee (IACUC). A431 cells (ATCC, Manassas, Virginia) were implanted subcutaneously into the right flank via an injection of cells in 50% Matrigel (BD Biosciences, San Jose, California) and 50% Dubelco’s modified essential media (DMEM) [10% fetal bovine serum (FBS), Pen/Strep] media, and grown for 8 to 14 days until reaching a volume of , then an injection of a mixture of 0.1 nmol of a epidermal growth factor receptor (anti-EGFR) Affibody (Affibody, Solna, Sweden) labeled with IRDye 800CW and 0.1 nmol of a nontargeted imaging control Affibody labeled with IRDye 680RD were made. Animals were sacrificed 1-h postinjection, the tumors removed, imbedded in Tissue Tek® Optimum Cutting Temperature medium (Sakura Finetek USA Inc., Torrance, California) and flash frozen. Tissues were stored at until sectioning. The frozen tumors were cut into 10 μm sections using a Leica 1850 M cryostat and fluorescence images of the tissues were immediately obtained on the Odyssey Scanner at a resolution of 21 μm (LICOR Biosciences, Lincoln, Nebraska) for both targeted and untargeted tracers. After fluorescent imaging, the tissue sections were stained with H&E by the Research Pathology Services at Dartmouth. The H&E stained tissues were imaged at .
A schematic representation of key elements of the method is shown in Fig. 1. First, the intensity map of the nontargeted fluorescent signal was subtracted from that of the targeted Affibody, to normalize for regions of nonspecific tracer pooling. Then, the normalized fluorescent images were segmented into 256 intensity levels [Fig. 1(a)] and then aligned and overlaid upon the H&E images prepared from the same section [Figs. 1(b) and 1(c)]. Regions of interest were identified by visual inspection of the fluorescent contour maps [Fig. 1(d)], selecting regions showing a close spacing of contours. Axes parallel and perpendicular to the concentration gradient were identified [Figs. 1(e) and 1(f)] to apply directional Fourier analysis.
These regions of the H&E images were evaluated for directional anisotropy by azimuthally averaged frequency power spectrum analysis, using a variation of the method introduced in Ref. 5. The method of generating directional FSF spectra is as follows: after identification of the region of anisotropic tracer concentration, a square patch of the tissue section, centered on the point at which the two axes of anisotropy intersect [Fig. 2(a)], is selected and rotated so that the axis of interest was at [Fig. 2(b)]. The three-color image was converted to an intensity map, as shown in Fig. 2(c), and the intensity of the cell nuclei was set to zero. This had a dual effect: first, it reduced high-frequency contributions from nuclear structures that were unlikely to have an effect on tracer transport; and second, it ensured that spatial frequencies corresponding to gross spacing between cell nuclei would be below 5 cycles/cm, allowing elimination of these with a low-pass filter. The two-dimensional (2-D) FT of the rotated intensity map was then generated by the built-in MATLAB function FFT2.
The power spectrum of an intensity map , is written as ,10 where and are the real and imaginary parts of the FT, respectively. Following the method established in Ref. 4, we decompose the power spectrum into three components, so that , where is the characteristic frequency that defines similarity between image regions, is the contribution of the spectrum that is specific to the region, and is a noise component that comes from instrumentation artifacts and the analytical limitations of the finite Fourier transform. To extract in the direction of interest, the azimuthally averaged power spectrum of a 7-deg arc of a sub-region of the 2-D FT centered around the principal axis of analysis was generated by numerically integrating the signal at 1 deg steps. The background signal was generated by integrating over the remaining 352 deg of the 2-D FT, sub-region , in Fig. 2(d), again by numerical integration at steps of 1 deg. Because of the high variability in FT traces over the large range of , the expected contribution of and is zero, and by definition , givingFig. 1(f), up to six to eight FSF spectra were generated per region. Similarity between these noise-reduced traces was established by taking their inner product. Noting that, by definition, region-specific spectra are orthogonal to both characteristic spectra and other region-specific spectra, so that , and gives:
In Figs. 3(a) and 3(b), individual FSF spectra from the principal and secondary axes of analysis of a representative image are plotted on the same axis, and the geometric mean of the sample is overlaid on them. The natural variation expected from a biological sample can be seen, as well as a striking correlation between the individual traces and the mean. In Fig. 3(c), the power spectrum of the two directional spectra is plotted on the same axis, showing spatial frequencies unique to each direction, with very little overlap. The method was applied to several regions of interest derived from contour maps, without regard to any apparent similarity of the underlying H&E staining. Figure 4 shows representative images with similar structures of tracer distribution anisotropy, which overlie regions of tumor stroma that have a markedly different visual appearance. The specific frequencies that identify stromal structural elements are not identical between different regions, but the characteristic directional spectra in the direction of flow can be seen to be distinct from those in the direction perpendicular to flow in each case.
We show here that the direction of free tracer transport can be differentiated from that of impeded tracer transport by comparison of the different characteristic spatial frequencies of the underlying eosin stain pattern. The system we use, H&E histopathology, is one that is well established in evaluation of cancer in the clinic. The pink stain of eosin, typically associated with protein, provides the signal from which we extract the characteristic directional FSF spectrum. The evaluative potential of this method can be extended by developing a staining design strategy that minimizes extraneous information, for example, by eliminating the hematoxylin stain, or that adds information, such as by locating microvessels by staining for vascular components.
It should be noted that the individual frequencies contained in a power spectrum do not have a straightforward correlation to features in an image, except in the case of an image with a regularly repeating sinusoidal pattern. For instance, the presence of a cluster of many high spatial frequencies is not necessarily an indicator of a large number of small features. It would be a fundamental misapplication of the method to expect that the presence of specific frequencies is an indicator of high or low transport. The correlation between spatial frequencies and visual patterns forms a distinct field of optical research and is beyond the scope of this letter. An excellent introduction to the topic can be found in Ref. 11.
Although there is an intuitive sense of what “the same” means from a pathologist perspective, the question of how to define similarity and difference, is not specific to spectral analysis or even to the natural sciences and can be found underlying arguments as diverse as biological taxonomy and the literary theory of deconstruction.12 For the purpose of FSF analysis, the key to identifying a characteristic spectrum in a particular direction is to differentiate the signal from a suitable background signal. The orthogonal spectrum, because it uses the same background, can always be distinguished from the principal direction if it has a distinct spectrum.
Because the fluorescent overlays show tracer distribution to be anisotropic, and there is some indication of this anisotropy in the H&E images even before analysis, the results presented here are a quantitative confirmation of what can be surmised. There are two principal experimental extensions of this method. First, as shown here, it provides a robust quantitative characterization of spatial anisotropy in systems where it can already be qualitatively detected. Directional FSF analysis can also be applied to imaging systems where a spatial anisotropy is expected but difficult to detect. Such would be the case in two-photon (2P) imaging of biological tissue, in which a second harmonic of the emission frequency is generated by collagen fibrils, and a variation in signal can be expected depending on the ordered state of the collagen [reviewed in Ref. (13)]. One distinct advantage of this application of 2P imaging is that it allows for real-time monitoring of drug distribution, while allowing for simultaneous capture of the signal used for FSF analysis. Likewise, it may be possible to monitor real-time destruction of transport barriers by activable nanoparticles14 designed for that purpose. A further extension of the method can be made to a 3-D case by stacking 2-D image slices.
In principal, the method of directional FSF analysis can be applied to any signal or imaging system. But there are two prerequisites for such a system. First, the signal to be analyzed must be available as a spatially distributed intensity map in which the intensity signal has some (known or unknown) relationship to the question at hand. Without this format, an informationally relevant FSF spectrum cannot be generated. Second, a suitable method must be identified for generating a background signal against which to isolate the signal of interest. This is especially important for biological applications, in which background noise can be on the same order of magnitude as the information.
In many tissues, structural anisotropy is strongly correlated with functional parameters. Diffusion tensor MRI can identify tracts of bundled neurons15 and muscle fibers,16 by the preferential diffusion of water along them, and automated methods have been explored.17 MRI tractography is possible because of the physics of the detection system, which allows for a large number of measurement directions (typically around 20) and the acquisition of a large number of data sets (typically thousands). An optical analogue of MRI tractography would be of significant benefit in the detection of anisotropic drug delivery in tumor. A key component will be an automated quantitative method of identifying similarities and differences between adjacent flow regions in optical maps, which we have demonstrated here. In future work, we hope to apply this method to images stained for microvasculature, to allow a discussion of fluid flow as it pertains to barriers to drug distribution. In addition, we will generate an automated scanning protocol that will examine steps along the track of the direction of interest to detect paths that have similar spectral characteristics that branch from or extend the initial path.
The current research was supported in part by the Professional Staff Congress/CUNY Professional Development Fund (SR), and the Institute for Ultrafast Lasers and Spectroscopy (RRA), NIH grants U54CA151662 (BWP, KSS, PJH) and R01CA156177 (BWP, KSS). The authors would like to thank Professors David S. Rumschitzki, PhD and Jeffrey Morris, PhD, of CCNY for providing research space, and thoughtful discussion, and Mary C. Schwab, lead technologist at Dartmouth Translational Research Shared Resources for H&E staining and imaging.
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Stewart Russell is a visiting research scientist at Dartmouth College, and holds posts in the mechanical engineering, and the master’s program in management information systems at the City College of New York (CCNY). He has a PhD in biomedical engineering, and was co-PI on the Bayer/CCNY Wireless Blood-glucose Monitor project. He holds patents on several fluorescence-based nanoparticle methods. His current focus is on laser scattering analysis methods for characterization of native and therapeutic nanoparticle transport.
Kimberley S. Samkoe is a research assistant professor in the Department of Surgery at Geisel School of Medicine and an adjunct assistant professor at Thayer School of Engineering at Dartmouth College. She obtained her PhD in biophysical chemistry from the Department of Chemistry at the University of Calgary, Alberta, Canada, in 2007. Her current research involves fluorescence imaging of targeted agents for cancer diagnosis and therapeutic monitoring, as well as photodynamic therapy.
P. Jack Hoopes is a professor of surgery, radiation oncology and biomedical engineering, with specialty training in veterinary pathology (DVM) and radiation biology/oncology (PhD), and specialization in development and use of animal models for translational biology/oncology and imaging research (MRI, CT, ultrasound, PET, fluoroscopy, and fluorescence imaging) for large and small animal models. He is director of Surgical Research Laboratories and the Center for Comparative Medicine and Research. Project PI–NCI funded Dartmouth Center for Cancer Nanotechnology Excellence.
Brian W. Pogue is professor of engineering, physics and astronomy, and surgery at Dartmouth College in Hanover, New Hampshire, USA. He has a PhD in medical/nuclear physics from McMaster University, Canada. He works in the area of optics in medicine, with a focus on novel imaging systems for characterizing cancer and imaging therapy. He has published over 230 peer-reviewed papers in the areas of biomedical optics, and his research is funded by the National Cancer Institute.