**1. Introduction**

Breast cancer is a significant health problem not only in the United States but globally and was the second leading cause of cancer-related death in the United States in 2018 [1]. Mammography, MRI (Magnetic Resonance Imaging), and B-mode ultrasound are the three most common imaging modalities used for breast cancer screening [2,3]. However, each of these modalities has its own unique shortcomings. The sensitivity of mammography in detecting breast lesions decreases in women with high-density breast tissue, and high-density breasts are considered to be more at risk for developing breast cancer [3,4]. In dense breasts, MRI can be used in conjunction with breast mammography to detect breast tumors [5,6]. Nevertheless, the operational cost and availability of MRI imaging limit the

accessibility of this modality. Conventional B-mode ultrasound (US) is a high-sensitivity, non-ionizing, and low-cost tool that is widely used for screening various types of human tissues [7,8]. However, false positives due to ultrasound screening result in many unnecessary biopsies [9,10]. Therefore, a more effective breast cancer screening tool is sought.

Photoacoustic tomography (PAT) is an imaging methodology that uses light absorption by endogenous or exogenous chromophores, and subsequent US pressure wave generation for imaging. Photoacoustic (PA) imaging has been demonstrated to be useful for a variety of medical and biological diagnostic applications, including early cancer detection [11–14]. Biomarkers such as vascularity and hypoxia have been shown to have diagnostic value in the di fferential diagnosis of various types of cancers including breast cancer [15–17]. In addition, when PA is augmented with nano-sized contrast agents, it can provide a reliable platform for the molecular imaging of cancer and its sub-types [18–21]. In clinical applications, PA imaging has been shown to produce real-time molecular and functional information with high resolution at relevant depths [19,22].

Several PAT imaging systems with di fferent illumination and acquisition modes have been developed for breast cancer imaging. However, the observed limitations of these systems are in part due to the non-optimum acoustic signal acquisition or illumination methodologies used. Our presented system is meant to be non-invasive (i.e., both illumination and acquisition are external), with the illumination and measurement system external to the imaged tissue. Our method images a cross-section inside the cylindrical US transducer array by illuminating the targeted area using a ring beam. Therefore, the light has to only di ffuse half the tissue diameter that is encountered when using side illumination. Point or di ffuse illuminations are suitable for imaging cross-sections close to the point of light entry [11,23–27], and the given fluence drops with light propagating through the tissue towards higher vertical depths as shown in Figure 1. This could make it di fficult to access areas close to the chest well.

**Figure 1.** The three methods of illumination for photoacoustic tomography (PAT) imaging that are compared in this study, with the definitions of vertical and cross-sectional imaging depths.

Other PAT imaging methods, such as the PA mammoscopy system [28], compress the tissue for better light penetration but can cause discomfort, or a loss of important PA biomarkers arises from the presence of blood by pushing the blood out of the tissue. One type of full-ring illumination system uses an acoustically penetrable optical reflector (APOR). However, APORs can only support low laser energies and US transmission through the reflector is highly angle-dependent [29–31]. Other illumination methods for deep tissue illumination include internal illuminations [32,33]. However, internal illumination is di fficult to develop for breast imaging applications. Therefore, it is vital to develop an alternative solution for improving the uniformity of energy distribution within the breast tissue for more accurate PAT imaging.

Ultrasound tomography (UST) using a ring-shaped US transducer has shown promising results in breast cancer screening [34–38]. In this work, PAT imaging is combined with this novel full ring UST system. The PA imaging modality can be easily combined with UST since both modalities share the same acquisition hardware. For this reason, the addition of the PAT to the UST is straightforward and will provide valuable functional information about a given tissue and is expected to improve the diagnostic capability of breast US for physicians.

The design of our combined UST/PAT imaging system has been previously presented [39–42]. This setup uses a ring illumination in conjunction with a ring US transducer for combined UST/PAT imaging. The ring-shaped beam in this system is generated by using a cone mirror and a parabolic reflector. This work specifically compares three di fferent illumination methodologies for PAT imaging: full-ring, di ffuse, and point illuminations. Using new findings from the three methods, it aims to show that full-ring illumination is the most e ffective method for creating PAT images due to its inherent cross-sectional fluence uniformity across vertical imaging depths (Figure 1). This is especially important for breast cancer screening when imaging close to the chest wall proves di fficult.

The three illumination methods are compared by imaging a three-layer polyvinyl chloride (PVC) tissue-mimicking phantom to gauge the advantages and disadvantages of common PAT imaging techniques. The experiments presented in this paper also all use the same data acquisition system and settings. Comparisons are made between PAT amplitudes for each cross-section and illumination methodology. Furthermore, the optimum position of the ring beam with respect to the targeted cross-section is examined.

#### **2. Material and Methods**

#### *2.1. UST*/*PAT Acquisition System*

A 200 mm diameter, 256-element ring US transducer (Analogic Corporation, Canada) with a center frequency of 2 MHz and bandwidth of 60% was used for all data acquisition. The presented system has a measured lateral resolution of 1 mm as determined by measuring a 200 micrometer light-absorbing string. This transducer has an element pitch of 2.45 mm and a height of 9 mm. The scattered US waves from a PA imaging event are recorded by all 256 elements using a sampling frequency of 8.33 MHz. As shown in Figure 2a, the US ring transducer is housed in an acrylic tank and is supplied with degassed, distilled water. During PAT imaging, the ring US transducer uses a 10 dB linear, time gain compensation (TGC) for acquiring the data, which is designed to optimize the signal-to-noise ratio (SNR) for the given phantom.

#### *2.2. Laser Source and Light Illumination Schemes*

A tunable, 10 nanoseconds pulsed laser (Phocus Core, Optotek, Carlsbad, CA, USA) was used for all PAT imaging experiments. This laser generates around 100 mJ per pulse at 532 nm. In the full-ring illumination mode, a large parabolic reflector (P19-0300, Optiforms Inc., Temecula, CA, USA) was used with a 10 mm diameter cone mirror (68-791, Edmund Optics, Barrington, NJ, USA) to create the 4 mm ring-shaped beam on the phantom surface (Figure 2a). Since the beam position is stationary, neither the cone mirror nor the parabolic reflector is mobile. The ring location was adjusted across each cross-section by translating the phantom in the vertical direction (Figure 2b). For the di ffused-beam experiments, a 120 grit ground glass di ffuser (DG10-120, Thorlabs Inc., Newton, NJ, USA) was placed in the laser light path inside the water tank after removing the cone mirror (Figure 2c). Finally, point illumination only uses the 45-degree mirror for directing the laser beam onto the phantom (Figure 2d).

**Figure 2.** (**a**) PAT experimental setup showing the water tank, ring ultrasound (US) transducer, and the translational stages. The experimental setups for the (**b**) full ring, (**c**) diffuse-beam, and (**d**) point illumination of the phantom.
