Neither binding nor internalization were observed in non-target cells, demonstrating absolute specificity and enhanced internalization of NPs by PSMA-expressing PC cells. Moreover, the accumulation of targeted NPs within LNCaP cells appeared as perinuclear punctate staining. This suggests internalization via receptor-mediated endocytosis and intracellular accumulation in endolysosomes [90].

For the characterization of active internalization by endocytosis, cells were incubated at two different temperatures: 4 ◦C and 37 ◦C. As endocytosis is an energy-dependent process, we assumed there will be higher accumulation at 37 ◦C, while at 4 ◦C only cell surface binding to the PSMA receptor may occur [90]. Confocal microscopy analysis (Figure 11) revealed that the uptake of targeted NPs was based on an endocytic process, with 17 red fluorescent dots per target LNCaP cell and an average intensity of 120 ± 60 (A.U.) at 37 ◦C compared to 4 dots per target cell and average intensity of 63 ± 30 (A.U.) at 4 ◦C. It should be noted that cellular processes other than endocytosis, such as diffusion, are also impeded at low temperatures and it is possible that TL binding to the PSMA receptor encouraged the diffusion of the NPs into the cell [91].

**Figure 11.** Characterization of active internalization of targeted NPs by LNCaP cells. Confocal laser microscopy images of LNCaP cells, following 1 h incubation with targeted NPs labeled with Cy7 (samples were diluted 1:400 (*v*/*v*) in FBS-free medium) at two different temperatures: (**A**) 37 ◦C; (**B**) 4 ◦C. Nuclear DNA staining achieved with Hoechst 33342 (1 μg/mL) is marked in blue, whereas NPs labeled with Cy7 are marked in red. Scale bars denote 20 μm.

#### *3.7. Selective Growth Inhibition of CTX-Loaded NPs In Vitro*

Cytotoxicity assays were performed using an XTT-based colorimetric cell proliferation kit with PSMA-positive LNCaP cells and compared to PSMA-negative PC cells (PC-3). It is evident from Figure 12 that CTX-loaded NPs displayed a remarkable dose-dependent growth inhibitory effect on target LNCaP cells which overexpress PSMA, with an outstanding half-maximal inhibitory concentration (IC50) of 4.0 ± 2.0 pM. In contrast, these NPs did not exhibit any substantial growth inhibitory effect on PC-3 which lacked PSMA expression.

Cells exposed to free CTX in a growth medium for 72 h served as control and IC50 values obtained with LNCaP and PC-3 cells were similar: 0.25 ± 0.04 nM and 0.18 ± 0.04 nM, respectively (Figure 13).

The ratio of 62.5-fold between the IC50 values of free CTX and CTX-loaded NPs towards target LNCaP cells can be explained by the different exposure times and the targeting ability of the CTX-loaded NPs. The active targeting of CTX-loaded NPs enables the accumulation of NPs in target PC cells and thus a more potent anti-tumor activity. As a result, even at shorter exposure times (24 h, compared to 72 h of free CTX), the cytotoxic effect of the NPs was more significant against LNCaP target cells, as evident from the lower IC50 value. Moreover, it has been reported that the IC50 of free CTX exposure for 48 h towards LNCaP cells is 2.6–3.1 nM, which supports this conclusion [92,93]. Another important aspect that was achieved in the current study is the remarkable LC of 168 ± 3 mg CTX/g SA-PEG. In this respect, based on the concentration of NPs that was determined by Nanosight, we calculated that each NP carries a cargo of as much as ~700,000 CTX drug molecules. This remarkably large amount of potent cytotoxic drug molecules that are efficiently and specifically internalized by target PC cells, not only eliminate these tumor cells but may also kill neighbor PC cells upon apoptotic and lytic cell death. It is important to emphasize that drug-free NLCs were found to be non-toxic to cells.

**Figure 12.** Selective growth inhibition of LNCaP target cells by CTX-loaded NPs. Cell growth inhibition as a function of the concentration of CTX encapsulated within NPs (0.6:1 CTX:SA-PEG molar ratio) in LNCaP target cells (circles) and non-target PC-3 cells (triangle). Values presented are means ± SE, n = 3 (three independent experiments, each performed in triplicates). The sigmoidal model curve of LNCaP cells was fitted using Equation (3).

**Figure 13.** Inhibitory effect of free CTX. Cell viability as a function of free CTX concentration of LNCaP target cells (circles) and PC-3 non-target cells (triangle). Values presented are means ± SE, n = 3 (three independent experiments, each performed in triplicates). Sigmoidal model curves were fitted using Equation (3).

#### **4. Conclusions**

Towards the efficacious treatment of PC, CRPC, and mCRPC, we herein developed novel selectively PC-targeted NPs harboring a cytotoxic drug cargo. These targeted NPs demonstrated efficient encapsulation of CTX, high specificity to, and potent eradication of, PSMA-positive target PC cells. Moreover, targeting NPs to the PSMA receptor resulted in uptake via receptor-mediated endocytosis, thus facilitating the intracellular release of the drug cargo, enabling the therapeutic activity. Another important aspect of selective internalization of these targeted NPs is the potential to overcome cancer MDR by evading MDR efflux pumps that extrude a broad spectrum of anticancer drugs from multidrugresistant cancer cells. This selectively targeted nano-delivery platform should be able to enhance the efficacy of PC treatment, while reducing drug doses and minimizing untoward toxicity. Hence, these novel findings bear promise towards the improvement of patient survival rates and quality of life. Future in vivo studies are warranted that should assess the therapeutic efficacy of the targeted NPs against human PC in a xenograft murine model.

**Author Contributions:** Investigation, L.C.; writing—original draft preparation, L.C.; writing—review and editing, Y.G.A. and Y.D.L.; supervision, Y.G.A. and Y.D.L. All authors have read and agreed to the published version of the manuscript.

**Funding:** This research received no external funding.

**Institutional Review Board Statement:** Not applicable.

**Informed Consent Statement:** Not applicable.

**Data Availability Statement:** The data presented in this study are available on request from the corresponding author.

**Acknowledgments:** We wish to thank Nitsan Dahan for his expert assistance with confocal laser microscopy. Moreover, we thank Moris Eisen, and Kseniya Kulbitski, Technion-Israel Institute of Technology, Haifa, Israel, for professional assistance with the SA-PEG-TL conjugation process. We also wish to thank O. Ben-Izhak from the department of Pathology at the Rambam Medical Center, Haifa, Israel for performing the immunohistochemistry analysis of PSMA expression.

**Conflicts of Interest:** The authors declare no conflict of interest.

#### **References**

## *Article* **Combined Antitumor Therapy Using In Situ Injectable Hydrogels Formulated with Albumin Nanoparticles Containing Indocyanine Green, Chlorin e6, and Perfluorocarbon in Hypoxic Tumors**

**Woo Tak Lee 1, Johyun Yoon 1, Sung Soo Kim 1, Hanju Kim 1, Nguyen Thi Nguyen 1, Xuan Thien Le 1, Eun Seong Lee 2, Kyung Taek Oh 3, Han-Gon Choi 4 and Yu Seok Youn 1,\***

**Abstract:** Combined therapy using photothermal and photodynamic treatments together with chemotherapeutic agents is considered one of the most synergistic treatment protocols to ablate hypoxic tumors. Herein, we sought to fabricate an in situ-injectable PEG hydrogel system having such multifunctional effects. This PEG hydrogel was prepared with (i) nabTM-technique-based paclitaxel (PTX)-bound albumin nanoparticles with chlorin-e6 (Ce6)-conjugated bovine serum albumin (BSA-Ce6) and indocyanine green (ICG), named ICG/PTX/BSA-Ce6-NPs (~175 nm), and (ii) an albumin-stabilized perfluorocarbon (PFC) nano-emulsion (BSA-PFC-NEs; ~320 nm). This multifunctional PEG hydrogel induced moderate and severe hyperthermia (41−42 ◦C and >48 ◦C, respectively) at the target site under two different 808 nm laser irradiation protocols, and also induced efficient singlet oxygen (1O2) generation under 660 nm laser irradiation supplemented by oxygen produced by ultrasound-triggered PFC. Due to such multifunctionality, our PEG hydrogel formula displayed significantly enhanced killing of three-dimensional 4T1 cell spheroids and also suppressed the growth of xenografted 4T1 cell tumors in mice (tumor volume: 47.7 ± 11.6 and 63.4 ± 13.0 mm3 for photothermal and photodynamic treatment, respectively, vs. PBS group (805.9 ± 138.5 mm3), presumably based on sufficient generation of moderate heat as well as 1O2/O2 even under hypoxic conditions. Our PEG hydrogel formula also showed excellent hyperthermal efficacy (>50 ◦C), ablating the 4T1 tumors when the irradiation duration was extended and output intensity was increased. We expect that our multifunctional PEG hydrogel formula will become a prototype for ablation of otherwise poorly responsive hypoxic tumors.

**Keywords:** photothermal therapy; photodynamic therapy; combined antitumor effect; oxygenation; hydrogel; hypoxic tumor

#### **1. Introduction**

Photodynamic therapy (PDT) has become a prominent therapeutic modality for cancer treatment [1–3]. Briefly, PDT is based on the clinical use of reactive oxygen species (ROS) such as singlet oxygen (1O2) or hydroxyl radicals (OH) generated from oxygen (O2) using photosensitizers (PS) and light energy. The intracellularly generated ROS are able to damage cancer cells and tumor vasculature and also induce immune-mediated tumor

**Citation:** Lee, W.T.; Yoon, J.; Kim, S.S.; Kim, H.; Nguyen, N.T.; Le, X.T.; Lee, E.S.; Oh, K.T.; Choi, H.-G.; Youn, Y.S. Combined Antitumor Therapy Using In Situ Injectable Hydrogels Formulated with Albumin Nanoparticles Containing Indocyanine Green, Chlorin e6, and Perfluorocarbon in Hypoxic Tumors. *Pharmaceutics* **2022**, *14*, 148. https://doi.org/10.3390/ pharmaceutics14010148

Academic Editor: Jun Dai

Received: 29 November 2021 Accepted: 6 January 2022 Published: 8 January 2022

**Publisher's Note:** MDPI stays neutral with regard to jurisdictional claims in published maps and institutional affiliations.

**Copyright:** © 2022 by the authors. Licensee MDPI, Basel, Switzerland. This article is an open access article distributed under the terms and conditions of the Creative Commons Attribution (CC BY) license (https:// creativecommons.org/licenses/by/ 4.0/).

suppression [3]. To date, several PDT agents, such as Photofrin®, Foscan®, and Visudyne®, etc., have been approved by the U.S. FDA, and the therapeutic availability of PDT has expanded toward cancer therapy. However, the clinical application of PDT is often restricted because most PS agents are effective only under focused laser illumination (650–900 nm) in the presence of sufficient oxygen. Unlike normal tissues, oxygenation in primary tumors is not regulated, and thus the partial oxygen pressure of tumors is often heterogeneous and can be remarkably low [4]. Interestingly, the O2 concentration within human tumors is known to be 1.3%~3.9% (mostly less than 0.33%), in comparison with 3.1%~8.7% within normal tissues [4,5]. The efficacy of many PDT and chemotherapeutic agents is seriously impaired by low oxygen concentrations within hypoxic tumors, which causes hypoxiainduced resistance of tumors to antitumor therapy [6]. Therefore, increasing the oxygen concentration in hypoxic tumors can be an effective method to overcome hypoxia-induced resistance of tumors to PDT [7–9].

Photothermal therapy (PTT) is also a promising antitumor modality that acts by raising tissue temperature [10]. Cancer cells are susceptible to temperature elevation and are often killed as a result of irreversible damage and subsequent induction of apoptosis even after exposure to temperatures >48 ◦C for only 5 min [11]. Many organic PTT agents emit heat in response to focused near-infrared (NIR) laser irradiation at wavelengths of 650−900 nm [10,12]. Because the hyperthermia can be concentrated within the area targeted by the NIR irradiation, PTT may create relatively little damage to surrounding healthy tissues [10,13]. Moreover, PTT can also be used to induce moderate local hyperthermia (39–42 ◦C) that promotes the delivery of oxygen as well as chemotherapeutics into tumor tissues due to elevated tumor blood flow [14,15]. This moderate PTT is useful in clinical settings because it attenuates the tumor hypoxic state and improves the efficacy of chemotherapeutic agents [16,17].

Albumin is a versatile carrier due to its excellent pharmaceutical advantages, such as biodegradability, biocompatibility, high stability, and solubility [18,19]. It has been used as a building material to fabricate a great number of therapeutic and diagnostic drug delivery systems, including nanoconjugates [20,21], nano-/micro-particles [22–25], and hydrogels of various sizes [26–29]. In practice, the nabTM (nanoparticle albumin-bound)-based paclitaxel formulation Abraxane® (Celgene Corp.) has displayed enhanced antitumor efficacy over the conventional paclitaxel formulation of Taxol®. In particular, this prototypical nabTM formula has sufficient flexibility and expandability to entrap various hydrophobic drugs, is better able to target the tumor gp60-mediated transcytosis pathway [18], and has been extensively used as a biocompatible component of injectable hydrogels [26–29].

In this study, we developed a multifunctional in situ-forming PEG-albumin hydrogel that was able to generate oxygen (O2) and singlet oxygen (1O2) and emit moderate or severe heat for the combination treatment of hypoxic tumors using chemotherapeutic, photodynamic, and photothermal therapy. This PEG hydrogel formula was designed to include (i) albumin nanoparticles containing paclitaxel (PTX), chlorin e6 (Ce6), and indocyanine green (ICG), and (ii) an albumin-stabilized perfluorocarbon (PFC) nano-emulsion (NE). The capability of this PEG hydrogel formula to generate heat and oxygen/singlet oxygen was carefully evaluated. The physicochemical and photo-induced properties of this multifunctional PEG hydrogel were also assessed by using relevant analyses. Finally, the combined chemotherapeutic, photodynamic and photothermal effects of this PEG hydrogel formula were carefully evaluated in 4T1 cell spheroids under hypoxic conditions and in 4T1 tumor-bearing mice.

#### **2. Materials and Methods**

#### *2.1. Materials*

Paclitaxel (PTX) was obtained from JW Pharmaceutical Corporation (Dangjin, South Korea). Indocyanine green (ICG) and icosafluoro-15-crown-5-ether (PFC) were purchased from Tokyo Chemical Industry Co., LTD. (Tokyo, Japan). Chlorin e6 (Ce6) was supplied by Frontier Scientific (Salt Lake City, UT, USA). Bovine serum albumin (BSA) and

2-iminothiolane (2-IT; Traut's reagent) were purchased from Sigma-Aldrich (St. Louis, MO, USA). Four-arm polyethylene glycol maleimide (4-arm PEG-MAL; Mw 20 kDa) was purchased from NOF Corporation (Tokyo, Japan). The 4T1 breast cancer cells were obtained from the American Type Culture Collection (ATCC; Rockville, MD, USA). LIVE/DEAD™ viability/cytotoxicity assay kits, Singlet Oxygen Sensor Green reagent (SOSG), and Cell-ROX™ Deep Red reagent were purchased from Thermo Fisher Scientific (Waltham, MA, USA). Anti-HIF-1α primary and Alexa Fluor®-488-conjugated goat anti-rabbit secondary antibodies were purchased from Abcam (Cambridge, MA, USA). Trypsin-EDTA and penicillin-streptomycin (P/S) solution were purchased from Corning (Somerville, NY, USA). Dulbecco's Modified Eagle Medium (DMEM) and fetal bovine serum (FBS) were purchased from Capricorn (Ebsdorfergrund, Hesse, Germany). All other reagents were obtained from Sigma-Aldrich unless otherwise indicated.

#### *2.2. Animals*

The animals were cared for in accordance with the Guide for the Care and Use of Laboratory Animals published by the United States National Institutes of Health. The protocols were also approved by the Institutional Animal Care and Use Committee (IACUC) of Sungkyunkwan University (IACUC number: 202106291, approval date 15 July 2021). The BALB/c *nu*/*nu* mice (female, 6 weeks old) were purchased from ORIENT BIO (Seongnam, South Korea). Mice were housed in microisolator cages on individually ventilated cage racks with ad libitum access to an autoclaved standard rodent diet (LabDiet 5008, Purina, St. Louis, MO, USA) and were kept under a 12 h light/dark cycle.

#### *2.3. Synthesis of Ce6-Conjugated BSA (BSA-Ce6)*

The BSA was covalently modified with Ce6 using a slight modification of the previous method [2]. Briefly, Ce6 (1 mmol), N-dicyclohexylcarbodiimide (DCC, 4 mmol), and Nhydroxysuccinimide (NHS, 5 mmol) were dissolved in 10 mL anhydrous dimethyl sulfoxide (DMSO) in a glass tube. Triethylamine (TEA, 4 mmol) was added to the mixture, and the reaction was performed in the dark at ambient temperature for 24 h. After removing the precipitate, the NHS-activated Ce6 was stored at −70 ◦C until required. Separately, 1.5 mL of Ce6-NHS (150 μmol) in DMSO was mixed dropwise with 50 mL of BSA (7.5 μmol; 0.1 M sodium borate buffer, pH 8.5) at a feeding ratio of 20:1 (Ce6:BSA). The reaction was allowed to continue at 450 rpm in the dark condition for 24 h. Unreacted Ce6 and DMSO were removed by using a dialysis membrane (MwCO: 10 kDa; Spectrum Labs, Rancho Dominguez, CA, USA) against 60% ethanol and deionized water (DW) for 48 and 24 h, respectively 24 h. Finally, Ce6-BSA was concentrated in DW using a centrifugal concentrator (MwCO: 30 kDa, Amicon® Ultra, Millipore, Burlington, MA, USA) and then lyophilized and stored at −20 ◦C for further use.

#### *2.4. Preparation of ICG/PTX/BSA-Ce6-NPs*

The ICG/PTX/BSA-Ce6-NPs were prepared as previously described using the nanoparticle albumin-bound (nabTM) technology with some adjustments [2,30–34]. In brief, 45 mg BSA, 5 mg BSA-Ce6, and 0.75 mg ICG were dissolved in 5 mL DW. Then, 5 mg PTX was dissolved in 0.1 mL of a 9:1 solution (chloroform:ethanol). These two solutions were gently shaken and homogenized by using a high-speed homogenizer (WiseTis® HG-15D: DAI-HAN Scientific Co., Seoul, South Korea) at 14,500 rpm for 3 min, and then passed through a high-pressure homogenizer (EmulsiFlex-B15 device, Avestin, Ottawa, ON, Canada) for nine cycles at 20,000 psi. The resulting dispersion was rotary evaporated for 15 min under reduced pressure at 40 ◦C to remove chloroform and ethanol. The nanoparticle fraction was centrifuged at 6000 rpm, and the supernatant was collected. The resulting suspension was lyophilized and stored at −20 ◦C until required.

#### *2.5. Fabrication of BSA-Stabilized PFC Nano-Emulsion*

The BSA-PFC nano-emulsion (BSA-PFC-NEs) was fabricated using a slight modification of the previous method [35]. A 300 μL aliquot of icosafluoro-15-crown-5-ether (PFC) was mixed with 4 mL of PBS containing 40 mg of BSA. The mixture was held in an ice bath and ultrasonicated (operation 8 s, interval 2 s) at an amplitude of 20% for 400 s. The resulting nano-emulsion was centrifuged (8000 rpm, 3 min), and the resulting nano-emulsion pellets were re-dispersed in 0.1 mL PBS.

#### *2.6. Characterization of ICG/PTX/BSA-Ce6-NPs and BSA-PFC-NEs*

The particle size and zeta potential of ICG/PTX/BSA-Ce6-NPs and BSA-PFC-NEs (DW) were measured by using a Zetasizer Nano ZS90 instrument (Malvern Instruments, Worcestershire, U.K.) at a dynamic light scattering mode. The surface morphology of the NPs was observed by transmission electron microscopy (TEM: JEM-3010, JEOL, Tokyo, Japan) and field-emission scanning electron microscopy (FE-SEM: JSM7000F, JEOL, Tokyo, Japan). The particle sizes of ICG/PTX/BSA-Ce6-NPs and BSA-PFC-NEs were measured at 0, 1, 2, 4, 8, 12, 24, 48, and 72 h at room temperature. In addition, UV-VIS-NIR spectral scans of ICG/PTX/BSA-Ce6-NPs or free ICG/Ce6 were also recorded over a range of wavelengths (300–900 nm) using a Synergy™ NEO microplate reader (Bio Tek, Winooski, VT, USA).

#### *2.7. Preparation of In Situ-Gelling PEG Hydrogel with ICG/PTX/BSA-Ce6-NPs and BSA-PFC-NEs*

The PEG hydrogel with ICG/PTX/BSA-Ce6-NPs and BSA-PFC-NEs (ICG/PTX/BSA-Ce6-NPs~PFC-NEs@Gel) was formulated using an in situ-gelling PEG hydrogel according to the protocol described previously with modifications [26,27,36]. To synthesize thiolated BSA (BSA-SH), BSA (500 mg) was dissolved in 1 mL of 100 mM PBS buffer (pH 8.0) containing 2-iminothiolane (2-IT, 12.5 mg), and the mixture was allowed to react by stirring for 2.5 h. The unreacted 2-IT was removed using a centrifugal concentrator (MwCO: 10 kDa, Amicon® Ultra, Millipore). The formula (1 mL) of the in situ-gelling PEG hydrogel comprised (i) ICG/PTX/BSA-Ce6-NPs (300 μL; 30 mg/mL), (ii) BSA-PFC-NEs (100 μL as a final emulsion), (iii) BSA-SH (150 μL; 500 mg/mL), and (iv) 4-arm PEG-MAL (450 μL; 11.4 mg/mL). Each component was mixed immediately, and gelation was optimized by monitoring the gelation time according to the molar ratio of 4-arm PEG-MAL and BSA-SH (BSA-SH:4-arm PEG-MAL = 1:3.3~4.5). The material was considered to be in a gel state if it did not flow when inverted.

#### *2.8. Characterization of ICG/PTX/BSA-Ce6-NPs~PFC-NEs@Gel*

The different PEG hydrogel formulas (Z1~Z6) with ICG/PTX/BSA-Ce6-NPs and BSA-PFC-NEs were designated as follows: (i) Z1 = PEG hydrogel; (ii) Z2 = PTX/BSA-NPs@Gel; (iii) Z3 = BSA-PFC-NEs@Gel; (iv) Z4 = PTX/BSA-Ce6-NPs@Gel; (v) Z5 = ICG/PTX-BSA-NPs@Gel; (vi) Z6 = ICG/PTX/BSA-Ce6-NPs~PFC-NEs@Gel (Table 1). General features of the hydrogel were characterized using photographs, photothermal images (FLIR E85 photothermal camera, FLIR Systems, Inc., Wilsonville, OR, USA), and fluorescence images (FOBI in vivo imaging system, NeoScience, Suwon, Korea).

#### *2.9. In Vitro and In Vivo Photothermal Imaging and Hyperthermia Monitoring*

A thermal imaging camera was used to measure the photothermal effect of the ICG/PTX/BSA-Ce6-NPs in vitro and in vivo using a slight modification of the previous method [26]. Naïve PEG hydrogel or ICG/PTX/BSA-Ce6-NPs@Gel in 1.5 mL Eppendorf tubes was irradiated for 5 min with an 808-nm laser (0.8 or 1.2 W/cm2 to induce mild or severe hyperthermia, respectively). To obtain photothermal images in vivo, naive PEG hydrogel or ICG/PTX/BSA-Ce6-NPs@Gel was injected into the tumors of mice, and the tumor regions were either irradiated or not irradiated with 808 nm laser (0.6–0.8 W/cm2) for 20 min. Temperature changes within the tumor areas were recorded and observed using a FLIR E85 photothermal camera (FLIR Systems, Inc., Wilsonville, USA).

#### *2.10. Singlet Oxygen Generation by ICG/PTX/BSA-Ce6-NPs*

Singlet Oxygen Sensor Green (SOSG) reagent was used to assess singlet oxygen generation from ICG/PTX/BSA-Ce6-NPs. Free Ce6 and ICG/PTX/BSA-Ce6-NPs were mixed with the SOSG reagent (final concentration 10 μM) in a 24-well plate. A 660-nm laser (10 mW/cm2) irradiated the respective sample wells for free Ce6 and ICG/PTX/BSA-Ce6-NPs under either normoxic or hypoxic conditions. Hypoxia was precisely achieved by incubating the hypoxic chamber in a hypoxic gas stream (5% CO2:95% N2 on a volume basis). While irradiating the sample wells with the laser for 60 min, a 100 μL aliquot of each sample was collected every 10 min. The intensity of the green fluorescence emitted from SOSG induced by 1O2 generation was determined using a multi-mode microplate reader (excitation 485 nm; emission 538 nm).

#### *2.11. Encapsulation Efficiency and Release Profiles of PTX, ICG, and BSA-Ce6*

The encapsulation efficiency (EE) of PTX, BSA-Ce6, and ICG into ICG/PTX/BSA-Ce6- NPs was examined. Briefly, 5 mg of the lyophilized ICG/PTX/BSA-Ce6-NPs was dissolved in 0.5 mL DW, and then 4.5 mL acetonitrile (ACN) was added to the NPs solution, followed by sonication for 30 min and centrifugation for 15 min (14,500 rpm). Subsequently, the supernatant was withdrawn for the quantification of PTX and ICG. The EE of BSA-Ce6 was calculated by measuring the supernatant concentration of BSA-Ce6 after the nanoparticle preparation. For the release test, a 5 mL ICG/PTX/BSA-Ce6-NPs solution (10 mM PBS, pH 7.4) was loaded into the PEG hydrogel. The drug release from the hydrogel was carried out using a gently rotated bench-top rotator (150 rpm, 37 ◦C). At predetermined times, a 0.15 mL aliquot was withdrawn and centrifuged (14,500 rpm, 10 min), and the fluorescence intensities of the supernatants were measured by using a fluorescence spectrometer (FluoroMax Plus; HORIBA Scientific, Tokyo, Japan). Three replicates of each sample were prepared and analyzed (n = 3). The excitation/emission wavelengths used for ICG and Ce6 were 420/526 and 405/670 nm, respectively. To determine the loading efficiency of PTX, the supernatant was subjected to reverse-phase high-performance liquid chromatography (RP-HPLC) using a PLRP-S Zorbax 100 RP-18 column (150 × 4.6 mm, 8 μm/300 Å; Agilent

Technologies, Palo Alto, CA, USA). An isocratic elution method was performed at a flow rate of 1.0 mL/min using a mixed solution (DW:ACN = 45:55). Eluates were monitored at 220 nm. Drug loading efficiency (%) was calculated as amount of drug entrapped/amount of drug loaded × 100.

#### *2.12. Cytotoxicity of ICG/PTX/BSA-Ce6-NPs in 4T1 Cell Spheroids Using LIVE/DEADTM Cell Assay*

Three-dimensional culture of 4T1 cells to form spheroids was performed by a slight modification of previously described methods [2,26,30]. LIVE/DEAD™ viability/cytotoxicity kits were used to visualize the death and viability of the 4T1 cells. Briefly, the 4T1 cells were cultured in DMEM media containing 10% (*v*/*v*) fetal bovine serum and 1% penicillin/streptomycin in a 5% CO2, 95% RH incubator at 37 ◦C. A 100 μL sample containing 4000 cells was seeded into each well of V-bottom cell plates (Shimadzu, Kyoto, Japan) designed to induce the formation of 3D multicellular spheroids. The 4T1 cell spheroids were then allowed to become established over three days before further experiments. The spheroids previously incubated with ICG/PTX/BSA-Ce6-NPs@Gel for 6 h were assigned to groups with or without laser irradiation (808 nm for 10 min or 660 nm for 60 min, respectively) and/or with or without an ultrasound (US) scan (US for 15 min) according to the previously described oxygen conditions (normoxic, hypoxic or oxygenated). The degree of hyperthermia induced was determined irradiation at 0.8 and 1.2 W/cm2 for mild and severe hyperthermia, respectively, where the target temperatures were set to 41–42 ◦C or >48 ◦C, respectively. The hypoxic condition was precisely achieved by incubating the hypoxic chamber in a hypoxic gas stream (5% CO2:95% N2 on a volume basis). The plates for samples were washed three times with DPBS, and the medium was replenished with fresh medium before laser or ultrasound treatment. After laser or ultrasound treatment, the cells were stained with Calcein-AM (for live cells) and Ethidium Homodimer-1 (for dead cells). Subsequently, the stained cells were observed by CLSM (excitation/emission wavelengths of 494/517 nm for Calcein-AM and 528/617 nm for Ethidium Homodimer-1, respectively).

#### *2.13. ROS Deep Assay at 4T1 Cell Spheroids*

The 4T1 cell spheroids were plated into each well of an 8-well chamber (Ibidi, Martinsried, Germany) and incubated with ICG/PTX/BSA-Ce6-NPs for 6 h. Cell spheroids were replenished with fresh DMEM (1% FBS) prior to laser irradiation (660 nm) and a US scan performed under normoxic, hypoxic, or oxygenated conditions. CellROX® Reagent was added to the cell spheroids at a final concentration of 20 μM and the plates were incubated at 37 ◦C for 2 h. Spheroids were washed three times with DPBS and stained with DAPI. The intracellular distribution of ROS in the spheroid cells was visualized using confocal laser scanning microscopy (CLSM; Carl Zeiss, Jena, Germany).

#### *2.14. HIF-1α Visualization at 4T1 Cell Spheroids*

The 4T1 cell spheroids were plated into each well of an 8-well chamber (Ibidi, Martinsried, Germany) and incubated with BSA-PFC-NEs for 1 h. The cell spheroids were subjected to a US scan under normoxic, hypoxic, or oxygenated conditions. Hypoxic conditions were achieved by incubating the hypoxic chamber in a hypoxic gas stream (5% CO2:95% N2 on a volumetric basis). The spheroids were fixed with 4% formaldehyde for 60 min and incubated for 90 min with 0.1% Triton X- 100, and then incubated with 1% BSA, 22.52 mg/mL glycine in PBST (PBS + 0.1% Tween 20) for 2 h to block unspecific binding of the antibodies. The resulting spheroids were further incubated with the diluted antibody solution (anti-HIF-1 alpha antibody-Alexa Fluor 488; 1/25) in 1% BSA in PBST in a humidified chamber for 1 h at room temperature. After removing the supernatant, the cell spheroids were washed three times in PBS, incubated with DAPI (DNA stain) for 4 h, and further rinsed with PBS. Samples were observed using CSLM.

#### *2.15. In Vivo Imaging of ICG/PTX/BSA-Ce6-NPs*

The in situ-gelling ICG/PTX/BSA-Ce6-NPs~PFC-NEs@Gel (100 μL) were directly injected into 4T1-cell-xenografted tumors on mice. At predetermined times, the fluorescence signals of Ce6 and ICG originating from the tumors were visualized at 630 and 730 nm, respectively, and at predetermined times using a FOBI in vivo imaging system (NeoScience, Suwon, Korea).

#### *2.16. In Vivo Photoacoustic Imaging*

Photoacoustic (PA) imaging was used to determine the saturation/distribution of oxygen or PFC gas inside the xenografted 4T1 tumors. When the tumors reached ~150 mm3 volumes, the mice were treated with (i) PBS; (ii) ICG/PTX/BSA-Ce6-NPs~PFC-NEs@Gel + US(+); (iii) ICG/PTX/BSA-Ce6-NPs~PFC-NEs@Gel + US(−). At before and after (30 min) intratumor injections, tumor oxygenation (as the oxyhemoglobin:HbO2 ratio) was visualized and recorded using a Vevo® LAZR-X Multimodal Imaging System (FujiFilm, VisualSonics Inc., Minato, Tokyo, Japan) set to the Oxy-hem mode (750 and 850 nm) at predetermined times.

#### *2.17. In Vivo Antitumor Efficacy in 4T1 Cell-Xenograft Mice*

The 4T1 cells (100 μL; 3 × 106 cells) were injected subcutaneously into the dorsal flanks of the mice. When the volume of the tumors reached ~150 mm3, the mice were randomly divided into five groups (G1−G5) as follows: (I) control: PBS; (II) chemotherapeutic = PTX/BSA-NPs@Gel (100 μg PTX (0.1 mL/mouse); (III) deoxy-photodynamic = ICG/PTX/BSA-Ce6-NPs@Gel + 660 nm (+); (IV) severe hyperthermia = ICG/PTX/BSA-Ce6-NPs@Gel + 808 nm(+); and (V) oxy-photodynamic plus moderate hyperthermia = ICG/PTX/BSA-Ce6-NPs~PFC-NEs@Gel + 660 nm (+) + 808 nm(+) + US(+) (Table 1). All treatments were injected intratumorally. Tumor volume and body weight of the mice were measured every day. Tumor volume (V) was calculated using the equation V = 0.5 × A × B2, where A is length (longest diameter) and B is width (shortest diameter). On day 14, the mice were sacrificed, and their tumors were harvested and photographed. The tumors were stored in formalin for further experiment.

#### *2.18. Data Analyses*

All data are presented as the mean ± standard deviation (SD). The significance of differences between treatment group results was determined using Student's *t*-test, and *p*-values < 0.05 were considered statistically significant.

#### **3. Results**

#### *3.1. Preparation and Characterization of ICG/PTX/BSA-Ce6-NPs and BSA-PFC-NEs*

The ICG/PTX/BSA-Ce6-NPs and BSA-PFC-NEs were fabricated using a high-pressure homogenizer based on the nabTM technology and ultrasonication, respectively (Figure 1). The sizes of the prepared ICG/PTX/BSA-Ce6-NPs and BSA-PFC-NEs were 175.8 ± 5.8 and 323.4 ± 6.6 nm, respectively (Figure 2A). Their zeta potentials were shown to be −40.3 ± 0.3 and −32.2 ± 1.2, respectively (Figure 2B). The ICG/PTX/BSA-Ce6-NPs were highly homogeneous and spherical, whereas the BSA-PFC-NEs appeared to be rather irregular and less homogeneous (Figure 2C). These solutions of ICG/PTX/BSA-Ce6-NPs and BSA-PFC-NEs showed the typical characteristics of a colloid or nano-emulsion (Figure 2D). In particular, the particle sizes of both the ICG/PTX/BSA-Ce6-NPs and BSA-PFC-NEs remained stable over a three-day evaluation period (Figure 2E). When observed at wavelengths ranging between 300 and 900 nm, the ICG/PTX/BSA-Ce6-NPs were found to have typical absorption profiles of indocyanine green and chlorin e6 of around ~800 nm and 660 nm, respectively. The Ce6-BSA, BSA-BC-NPs, and Ce6-BSA-BC-NPs were investigated (Figure 2F). In addition, the loading efficiency of PTX, ICG, and BSA-Ce6 was found to be ~73.0%, 53.2%, and 82.1%, respectively.

**Figure 1.** Schematic illustration of combined photothermal and photodynamic therapy for hypoxic tumors using ICG/PTX/BSA-Ce6-NPs~PFC-NEs@Gel under laser irradiation (660/808 nm) and ultrasound trigger.

**Figure 2.** (**A**) Histograms of particle sizes of ICG/PTX/BSA-Ce6-NPs and BSA-PFC-Nes. (**B**) Zeta potentials of ICG/PTX/BSA-Ce6-NPs and BSA-PFC-NEs. (**C**) TEM and FE-SEM images of ICG/PTX/BSA-Ce6-NPs (left: a) and BSA-PFC-NEs (right: b). (**D**) Photographs showing colloidal solution state for ICG/PTX/BSA-Ce6-NPs and BSA-PFC-NEs. (**E**) Physical stability of ICG/PTX/BSA-Ce6-NPs and BSA-PFC-NEs based on size changes over 3 days. (**F**) UV-VIS spectroscopic analyses for free ICG, free Ce6, and ICG/PTX/BSA-Ce6-NPs at 300–900 nm wavelengths.

#### *3.2. Preparation, Characterization, and Drug Release Profiles of ICG/PTX/BSA-Ce6-NPs~PFC-NEs@Gel*

Given that the gelation process depends on a rapid thiol-maleimide reaction, in situ gelling of the PEG hydrogel was evaluated by monitoring the relationship between the gelation time and the molar ratio of 4-arm PEG-MAL to BSA-SH. As shown in Figure 3A, reducing the 4-arm PEG-MAL to BSA-SH molar ratio sharply reduced the gelation time, with the PEG hydrogel becoming solid in approximately 1 min at a 4-arm PEG-MAL to BSA-SH ratio of 4.4:1.0, which was considered optimal to perform a series of related steps (i.e., material mixing, syringe filling, and injection). The in vitro release profiles of Ce6 (as a BSA-Ce6 conjugate), ICG, and PTX from ICG/PTX/BSA-Ce6-NPs~PFC-NEs@Gel were observed over 168 h. As shown in Figure 3B, >90% amount of BSA-Ce6 and ICG was released over 24 h with initial burst-out, and the remaining amount of both BSA-Ce6 and ICG appeared to be completely released from the PEG hydrogel after 144 h, which was appropriate for the in vivo laser irradiation protocol. In contrast, PTX displayed a gradual linear release pattern consisting without significant initial fast release until 168 h, which was presumably due to the hydrophobic property, tight binding to albumin molecules, and low solubility/diffusion of PTX. Thus, PTX seemed to have considerable time to be released from the PEG hydrogel. Subsequently, the hydrophobicity difference between PTX, ICG, and BSA-Ce6 led to a great discrepancy in the release pattern.

**Figure 3.** (**A**) Monitoring of gelation time for in situ-forming PEG hydrogel according to the 4-arm PEG-MAL:BSA-SH molar ratio. (**B**) Release profiles for BSA-Ce6, ICG, and PTX from ICG/PTX/BSA-Ce6-NPs@Gel.

Separately, the various hydrogel formulations (Z1−Z6) had distinct colors, ranging from transparent to milky, khaki green, or dark green, depending on ICG, Ce6, and PFC concentrations (Figure 4A). All formulas gelled easily within ~60 s and did not flow or drop when turned upside-down. As shown in Figure 4B,C, PEG hydrogels containing ICG (Z5 and Z6) produced significant hyperthermia, and those containing both ICG and Ce6 emitted strong fluorescence at the relevant emission wavelengths. The FE-SEM images showed that the PEG hydrogels contained considerable amounts of nanoparticles having sizes of ~150 and >300 nm, presumably corresponding to ICG/PTX/BSA-Ce6-NPs and BSA-PFC-NEs (Figure 4D).

**Figure 4.** (**A**) Photographs of PEG hydrogel formulation groups (Z1–Z6) in the regular (left) and upside-down (right) state. (**B**) Various images of PEG hydrogel formulation groups (Z1–Z6) (a) photograph at 10 min after gelation, (b) thermal images, (c) fluorescence images of ICG, (d) fluorescence images of Ce6. (**C**) Thermal images of glass vials containing the Z1−Z6 PEG hydrogel formulations during laser irradiation at 420/526 and 405/670 nm, excitation/emission wavelengths to identify ICG and Ce6, respectively. (**D**) FE-SEM images for the morphology of inner structures of ICG/PTX/BSA-Ce6-NPs@Gel.

#### *3.3. Photothermal and Photodynamic Activity of ICG/PTX/BSA-Ce6-NPs*

The induction of hyperthermia by ICG/PTX/BSA-Ce6-NPs@Gel when irradiated by the 808 nm laser was evaluated. As shown in Figure 5A,E, the temperature of the ICG/PTX/BSA-Ce6-NPs@Gel quickly increased from 26 to 56 ◦C within 5 min of initiation of 808-nm laser irradiation (1.2 W/cm2), whereas the temperature of the plain PEG hydrogel increased by only 3 ◦C. However, the temperature of ICG/PTX/BSA-Ce6-NPs@Gel increased to 42 ◦C in response to 808 nm laser irradiation (0.8 W/cm2) (Figure 5B). As shown in Figure 5C, when 808-nm laser irradiation (0.8 W/cm2) was applied intermittently in 3 min on/off cycles, the gel temperature reached ~41 ◦C after 30 s irradiation and returned to the basal temperature before additional irradiation. Separately, ICG/PTX/BSA-Ce6-NPs were clearly able to induce singlet oxygen generation, as evaluated by the fluorescence intensity caused by SOSG. Overall, the SOSG fluorescence appeared to increase with the duration of irradiation (Figure 5D). Singlet oxygen generation by ICG/PTX/BSA-Ce6-NPs appeared comparable to that of free Ce6, and much less singlet oxygen was generated under hypoxic conditions than under normoxic conditions. Furthermore, the tumor temperature

of 4T1 tumor-bearing mouse treated with ICG/PTX/BSA-Ce6-NPs@Gelplus 808-nm laser irradiation (0.8 W/cm2) maintained around ~41 ◦C for 20 min, whereas that of mouse treated with the naïve PEG hydrogel at the same condition did not significantly increase (Figure 5F).

**Figure 5.** (**A**) Temperature change profile for naïve PEG hydrogel and ICG/PTX/BSA-Ce6-NPs@Gel after targeted 808-nm laser irradiation (1.2 W/cm2) to induce severe hyperthermia with. (**B**) Temperature change profile for naïve PEG hydrogel and ICG/PTX/BSA-Ce6-NPs@Gel with or without targeted 808-nm laser irradiation (0.8 W/cm2) to induce moderate hyperthermia. (**C**) Temperature adjustment profile for ICG/PTX/BSA-Ce6-NPs@Gel during alternating on and off 808-nm laser irradiation (0.8 W/cm2) for 3 min cycles. (**D**) SOSG fluorescence profiles of Ce6 and ICG/PTX/BSA-Ce6-NPs in response to continuous 660 nm-irradiation (10 mW/cm2) as assessed by singlet oxygen generation under normoxic and hypoxic conditions. (**E**) Thermal images of (a) naïve PEG hydrogel and (b) ICG/PTX/BSA-Ce6-NPs@Gel during 808 nm laser irradiation (1.2 W/cm2). (**F**) In vivo whole-body thermal images of mice bearing 4T1 tumor injected with (a) naïve PEG hydrogel or (b) ICG/PTX/BSA-Ce6-NPs@Gel plus 808-nm laser irradiation (0.8 W/cm2) over 20 min.

#### *3.4. Cytotoxicity Evaluation of 4T1 Cell Spheroids Based on LIVE/DEAD™ Assay*

With regard to photothermal effects, when irradiated with the 808 nm laser at 1.2 W/cm2 for 10 min, 4T1 cell spheroids appeared to be absolutely dead, showing clear red fluorescence color, whereas 4T1 spheroids were almost all alive after irradiation with the 808 nm

laser at 0.8 W/cm2 for 10 min, showing mostly green fluorescence color (Figure 6A). With regard to photodynamic effects, the 4T1 cell spheroids cultured with PBS and ICG/PTX/BSA-Ce6-NPs/BSA-PFC-NEs seemed all alive when held under hypoxic conditions, irrespective of 660 nm laser irradiation. However, the 4T1 cell spheroids cultured with ICG/PTX/BSA-Ce6-NPs and irradiated with the 660 nm laser for 60 min were obviously dead when held under normoxic conditions, whereas non-irradiated 4T1 cell spheroids cultured under the same condition showed negligible cytotoxicity. Furthermore, the 4T1 cell spheroids incubated with ICG/PTX/BSA-Ce6-NPs/BSA-PFC-NEs and irradiated with the 660 nm laser (60 min) and also subjected to a US scan (15 min) were mostly dead because oxygen release from the BSA-PFC-NEs was induced by ultrasound scan energy, whereas the 4T1 cell spheroids cultured under the same conditions but without the US scan clearly remained alive (Figure 6B).

**Figure 6.** (**A**) CLSM images for 4T1 cell spheroids incubated with ICG/PTX/BSA-Ce6-NPs@Gel under 808 nm laser irradiation specifically adjusted to induce regular, moderate, or severe temperature. (**B**) CLSM images for 4T1 cell spheroids incubated with ICG/PTX/BSA-Ce6-NPs@Gel with or without 660 nm laser irradiation under normoxic, hypoxic, or oxygenated conditions. (**C**) CLSM images for 4T1 cell spheroids incubated with PTX/BSA-NPs@Gel. Green and red colors represent live and dead cells, respectively.

#### *3.5. Visualization of Hypoxia in 4T1 Cell Spheroids*

Hypoxic conditions were precisely achieved at the hypoxic chamber in a hypoxic gas stream (5% CO2:95% N2 on a volumetric basis), and the resulting hypoxia was demonstrated by visualizing the hypoxia-inducible factor (HIF)-1α at 4T1 cell spheroids. As shown in Figure 7A, spheroids treated at hypoxia were strongly stained with anti-HIF-1α antibody, displaying green fluorescence. On the other hand, spheroids treated at normoxia was not stained with anti-HIF-1α antibody because HIF-1α was negligibly expressed. However, 4T1 cell spheroids incubated with ICG/PTX/BSA-Ce6-NPs/BSA-PFC-NEs that were not subjected to a US scan displayed strong green fluorescence when cultured under hypoxic conditions. In contrast, the same 4T1 cell spheroids oxygenated by a US scan did not show the typical hypoxic fluorescence signal.

**Figure 7.** (**A**) CLSM images of 4T1 cell spheroids incubated with ICG/PTX/BSA-Ce6-NPs~PFC-NEs@Gel with or without US guidance under normoxic or hypoxic conditions in order to assess consequent HIF-1α production. (**B**) CLSM images of 4T1 cell spheroids incubated with ICG/PTX/BSA-Ce6- NPs~PFC-NEs@Gel with or without 660 nm laser irradiation under normoxic or hypoxic conditions in order to assess consequent intracellular ROS (singlet oxygen) production. (**C**) CLSM images of 4T1 cell spheroids incubated with ICG/PTX/BSA-Ce6-NPs~PFC-NEs@Gel with or without US guidance and irradiated with a 660 nm laser under normoxic or hypoxic conditions in order to assess consequent intracellular ROS production.

#### *3.6. Visualization of Singlet Oxygen Generation in 4T1 Cell Spheroids*

The intracellular ROS levels/distributions were visualized in each group of 4T1 cell spheroids by CLSM. When cultured under hypoxic conditions, all test groups of 4T1 cells incubated with ICG/PTX/BSA-Ce6-NPs~PFC-NEs@Gel did not exhibit strong red signals indicating significant intracellular ROS production, as detected by the ROS-sensitive dye CellROX® Deep Red, regardless of the use of 660 nm laser irradiation (Figure 7B). In contrast, when cultured under normoxic conditions, all 4T1 cell spheroids incubated with the various hydrogels and irradiated with the 660 nm laser displayed strong red fluorescence signals, and this phenomenon appeared even when the spheroids were cultured under hypoxic conditions but oxygenated by a US scan of the PEG hydrogel samples (Figure 7C).

#### *3.7. In Vivo Tumor Localization of ICG/PTX/BSA-Ce6-NPs~PFC-NEs@Gel in 4T1 Tumor-Bearing Mice*

After intratumor injection of ICG/PTX/BSA-Ce6-NPs~PFC-NEs@Gel, ICG and Ce6 fluorescence were visualized in xenografted 4T1 tumor-bearing mice. The ICG and BSA-Ce6 from the PEG hydrogel were diffusely distributed within the tumors. The considerable fluorescence levels were maintained until 12 h, but their NIR-fluorescence signals gradually decreased and then faded out 72 h after intratumor injection (Figure 8).

**Figure 8.** In vivo visualization of 4T1 tumor-bearing mouse after intratumor injection of ICG/ PTX/BSA-Ce6-NPs~PFC-NEs@Gel based on Ce6 and ICG fluorescence, at 630 and 730 nm, respectively.

#### *3.8. Photoacoustic Imaging of 4T1 Cell-Xenograft Tumors*

The accumulation of oxygen (as HbO2) produced in 4T1 tumors xenografted onto mice was visualized using photoacoustic (PA) analyses with high resolution and deep imaging depth (Figure 9). Groups were designated as either the PBS control or ICG/PTX/BSA-Ce6-NPs~PFC-NEs@Gel (G5 formula) with or without a US scan. The PEG hydrogel group (G5) with US scan under mild heat (41−42 ◦C) clearly displayed the greatest oxygen distribution (red signal) throughout the tumors on PA imaging: the HbO2 percentage was measured as 25.0% to 42.7% at before and after US, respectively, compared with the negligible HbO2-level changes (around 23.1%~29.6%) for PBS control and G5 + US(−). This result demonstrated that our ICG/PTX/BSA-Ce6-NPs~PFC-NEs@Gel had a significant ability to supply oxygen to the hypoxic breast tumors.

**Figure 9.** Photoacoustic images of 4T1 cell tumors of mice before and after (at 20 min) intratumor injection of PBS or ICG/PTX/BSA-Ce6-NPs~PFC-NEs@Gel with or without US trigger.

#### *3.9. Antitumor Efficacy in 4T1 Tumor-Bearing Mice*

The antitumor effects of ICG/PTX/BSA-Ce6-NPs~PFC-NEs@Gel were evaluated in 4T1 tumor-bearing mice. The tumor volumes for each group of mice were measured for 14 days after treatment (Figure 10A): the final tumor volumes for the G1 to G5 groups were 805.9 ± 138.5, 565.5 ± 41.6, 424.2 ± 90.6, 47.7 ± 11.6, and 63.4 ± 13.0 mm3, respectively. The G4 formulation with severe heat-induced regression of almost the entire tumor in each xenografted mouse due to the hyperthermal effect. In addition, the G5 formulation (with oxygen provided by the US scan and mild heat induced by the 808 nm laser and singlet oxygen generated by 660 nm laser irradiation) noticeably suppressed the tumors as compared to the G3 group that did not receive oxygen induced by the US scan. The body weight of the mice in the five treatment groups was maintained without significant change over 14 days, indicating that all mice were well cared for without deleterious effects during the entire therapy period (Figure 10E).

177

**Figure 10.** In vivo antitumor efficacy of each PEG hydrogel formulation in five treatment groups of 4T1 tumor-bearing mice (G1−G5): (I) control = PBS; (II) chemotherapeutic = PTX/BSA-NPs@Gel (100 μg PTX (0.1 mL/mouse)); (III) deoxy-photodynamic = ICG/PTX/BSA-Ce6-NPs@Gel + 660 nm (+); (IV) severe hyperthermia = ICG/PTX/BSA-Ce6-NPs@Gel + 808 nm(+); (V) oxy-photodynamic plus moderate hyperthermia = ICG/PTX/BSA-Ce6-NPs@Gel + BSA-PFC-NEs + 660 nm (+) + 808 nm(+) + US(+). All treatments were injected intratumorally. (**A**) Tumor volumes in each group over 14 days. (**B**) Representative photographs of 4T1-tumor-bearing mice at the indicated days after treatment. (**C**) Photographs of tumors excised from each treatment group. (**D**) Body weight change of 4T1-tumor-bearing mice in the different treatment groups over 14 days post treatment.

#### **4. Discussion**

Combined anticancer treatment using different therapeutic modalities has attracted great attention as an effective way of suppressing malignant tumors [37]. In particular, co-therapy using PDT and PTT may accomplish synergistic efficacy based on different modes of anticancer treatment [26,38]. While PDT is a well-established technique that is commonly used clinically, is minimally invasive, and produces few side effects [39], PTT is used for two distinct purposes in practice: (i) independent treatment to suppress tumors by severe hyperthermia and (ii) supportive treatment to improve the delivery of chemotherapeutics and oxygen into tumors by increasing tumor blood flow due to moderate hyperthermia. Moreover, concurrent use of chemotherapeutics could accelerate tumor ablation [40]. To this end, we focused on enhancing hypoxia-induced tumor suppression by the use of intratumorally injectable PEG hydrogel formulations based on

combined effects of chemotherapeutic, photodynamic, and photothermal modalities plus an oxygen-generation system.

For photodynamic therapy, Ce6 was chosen because it is one of the most efficient and safe second-generation PSs that is approved by the FDA for clinical use. It is sensitive to the light of a wide range of wavelengths (red to NIR: 650−900 nm), which penetrates relatively deep into tumors (presumably hypoxic regions) [39]. However, Ce6 is difficult to formulate on account of its poor solubility in both water and chloroform/dichloromethane; hence, albumin conjugates of Ce6 have been used to surmount this difficulty [2,41]. On the other hand, PFC formulated in albumin-stabilized nano-emulsion can be used to potentiate the photodynamic activity of Ce6 in hypoxic tumors by providing oxygen. Several methods have been introduced to supply or generate extra-/intracellular oxygen (i.e., CaO2/catalase [7], MnO2/H2O2 [9,42], and a PFC oxygen shuttle/US [35,43]). Among these, PFC is chemically inert and has been explored as an artificial blood substitute due to its high affinity for oxygen. Indocyanine green (ICG) was considered the first option considered for photothermal therapy because it is a deep-penetrating NIR dye approved by the US FDA; in addition, ICG is multifunctional and is also used as a diagnostic and photodynamic agent [44]. However, ICG has several undesirable properties, including concentration-dependent aggregation, poor aqueous stability, and rapid renal elimination from the body, which limit its clinical use [45,46].

The primary aim of this study was to fabricate an in situ-injectable "one-pot" delivery system optimal for including Ce6, ICG, PFC, and PTX in an attempt to achieve concurrent photodynamic, photothermal, and chemotherapeutic effects as a result of improved oxygenation of hypoxic tumors. Previously, we developed a simple method to prepare an in situ-gelling albumin-cross-linked PEG hydrogel system. This hydrogel system comprised only two components (thiolated albumin and 4-arm PEG-maleimide (20 kDa)) and is easily gelled within ~60 s in situ [26,27,36]. If required, the gelation time could be easily adjusted from approximately 15 s to 5 min according to the intended purpose [27]. Likewise, as shown in Figures 3 and 4, all formulations of our ICG/PTX/BSA-Ce6-NPs~PFC-NEs@Gel appeared to be gelled well within 60 s and to have acceptable elastic gel characters: all PEG hydrogel formulations (Z1−Z6) did not flow or drop when held upside-down for 60 s. In clinical settings, this 60 sec period can be critical to allow medical doctors or nurses to complete a series of injection steps: (i) mixing/dissolving powder materials with sterile water for injection, (ii) filling a syringe with the resulting solution, and (iii) immediately injecting the hydrogel into relevant sites before gelation occurs. Likewise, the gelation time for our PEG hydrogels, including those with added ICG/PTX/BSA-Ce6-NPs and BSA-PFC-NEs, was flexibly controlled for 30−120 s depending upon the BSA-SH:4 arm PEG-MAL molar ratio (=1:4.3−4.5). The prepared PEG hydrogel was confirmed to contain ICG/PTX/BSA-Ce6-NPs and BSA-PFC-NEs and displayed moderate to severe hyperthermal and photodynamic effects as measured by a photothermal camera and SOSG fluorescence (Figure 5). In addition, formation of the PEG hydrogel was the result of a specific reaction between artificially driven thiol groups in BSA and maleimide groups of the 4-arm PEG-MAL, and thus the delay or interruption of the release of BSA-Ce6, ICG, or PTX/PFC due to conjugation with PEG frames could be theoretically excluded. Consequently, BSA-Ce6, ICG, and PTX were almost released within 24~36 h (Figure 3B), and oxygen release from BSA-PFC-NEs seemed to be triggered by ultrasound and contributed to the creation of normoxic conditions.

Albumin nanoparticles/nano-emulsions conjugated with Ce6, ICG, and PFC were used as nano-systems inside the PEG hydrogel. Owing to problems with the solubility and stability of Ce6 and ICG, nabTM-technique-based albumin nanoparticles were considered to be an optimal formulation to stably load both Ce6 and ICG into the hydrogels. Previously, our group has optimized albumin nanoparticle formulations for Ce6 and ICG [2,30]. Basically, the strongly hydrophobic chemotherapeutic agent PTX played a role in physically cross-linking albumin molecules, and Ce6-conjugated BSA (BSA-Ce6) was embedded within the resulting ICG/PTX/BSA-Ce6-NPs during their preparation. Despite high water

solubility, ICG is known to strongly bind to albumin and thus is highly encapsulated into albumin nanoparticles because the aromatic and heteroaromatic rings of ICG interact strongly with hydrophobic pockets within albumin molecules [45,47]. As shown in Figure 2A,C, these two photothermal and photodynamic agents were simultaneously incorporated into albumin nanoparticles of ~175 nm diameter, and the resulting ICG/PTX/BSA-Ce6-NPs were able to provide both activities in response to the laser irradiations (660 and 808 nm). On the other hand, liquid PFC was easily formed through ultrasonication as a nano-emulsion (BSA-PFC-NEs) stabilized by surrounding albumin molecules, showing ~320 nm emulsion size, and thus appeared to be loaded into the PEG hydrogel. Importantly, the entrapped ICG/PTX/BSA-Ce6-NPs and BSA-PFC-NEs appeared not to undergo sudden release from the PEG hydrogel because the hydrogel mesh formed by four-arm PEG (Mn 5−20 kDa) is thought to be ~5 to 15 nm as estimated using the Flory–Rehner equation or rubber elasticity theory [48]. However, ICG and BSA-Ce6 of relatively small size (<6 nm) may be released through the mesh pores and move (in)to target cancer cells.

Our PEG hydrogel formulated with added ICG/PTX/BSA-Ce6-NPs and BSA-PFC-NEs had significant ability to kill 4T1 cell spheroids due to both 808 nm laser-induced hyperthermia and 660 nm laser-induced photodynamic cell death, accompanied by oxygen release from US-triggered BSA-PFC-NEs. Whereas cell death induced by PTX alone seemed insignificant, the production of naïve oxygen and singlet oxygen from our PEG formulations was clearly observed in anti-HIF-1α and ROS deep assays, respectively (Figures 7 and 8). In addition, our ICG/PTX/BSA-Ce6-NPs~PFC-NEs@Gel unambiguously increased oxygen supply to the tumor region as shown in the photoacoustic imaging (Figure 9), and the PEG hydrogel was retained within the area of the injection site for more than three days. This indicated that additional laser irradiation or US scanning might result in much better tumor suppression/regression. Overall, the severe hyperthermal tumor ablation effect of our formula combined with the strengthened photodynamic activity due to oxygenation and moderate hyperthermia (41–42 ◦C) appeared to potentiate hypoxic tumor ablation in 4T1 tumor-bearing mice.

#### **5. Conclusions**

In summary, the in-situ injectable PEG hydrogel with added ICG/PTX/BSA-Ce6-NPs (~175 nm) and BSA-PFC-NEs (~320 nm) was easily and quickly (~60 s) fabricated and displayed photothermal, photodynamic, and oxygen-supplying activity. The hypoxic tumors of mice treated with our PEG hydrogel formula and with 808/660 nm laser irradiation plus ultrasound scanning were unambiguously suppressed in volume compared with those of all other groups. This remarkable tumor size reduction is attributable to the strengthened photodynamic activity due to oxygenation and moderate hyperthermia (41~42 ◦C) by the multifunctional effect of loaded Ce6, ICG, and PFC. In addition, our PEG hydrogel formula had sufficient hyperthermal activity to ablate the 4T1 tumors when the irradiation time and output intensity increased. We believe our multifunctional PEG hydrogel formula can be a prototype for ablating otherwise unresponsive hypoxic tumors in coordination with deep-penetrating outer stimuli.

**Author Contributions:** Conceptualization: Y.S.Y. and W.T.L.; Methodology: W.T.L. and S.S.K.; Software: W.T.L. and N.T.N.; Validation: J.Y. and W.T.L.; Formal analysis: H.K. and W.T.L.; Investigation: E.S.L. and K.T.O.; Resources: H.-G.C.; Data curation: W.T.L., X.T.L. and Y.S.Y.; Writing—original draft preparation: Y.S.Y.; Writing—review and editing: E.S.L., K.T.O., Y.S.Y. and H.-G.C.; Visualization: Y.S.Y.; Supervision: Y.S.Y.; Project administration: Y.S.Y.; Funding acquisition: Y.S.Y. All authors have read and agreed to the published version of the manuscript.

**Funding:** This work was supported by National Research Foundation of Korea (NRF) grants funded by the Korean government (MSIT) (no. NRF-2019R1A2C2085292 and no. NRF-2019R1A5A2027340).

**Institutional Review Board Statement:** The study was conducted according to the guidelines of the Declaration of Helsinki and approved by the Institutional Review Board (or Ethics Committee) of Sungkyunkwan University (approved number of 202106291, approval date 15 July 2021).

**Informed Consent Statement:** Not applicable.

**Data Availability Statement:** Not applicable.

**Conflicts of Interest:** The authors declare no conflict of interest.

#### **References**

**Artem A. Sizikov 1, Petr I. Nikitin 2,3,\* and Maxim P. Nikitin 1,4,\***

	- 3 National Research Nuclear University "MEPhI", 115409 Moscow, Russia

**Abstract:** Nanoparticle-based technologies are rapidly expanding into many areas of biomedicine and molecular science. The unique ability of magnetic nanoparticles to respond to the magnetic field makes them especially attractive for a number of in vivo applications including magnetofection. The magnetofection principle consists of the accumulation and retention of magnetic nanoparticles carrying nucleic acids in the area of magnetic field application. The method is highly promising as a clinically efficient tool for gene delivery in vivo. However, the data on in vivo magnetofection are often only descriptive or poorly studied, insufficiently systematized, and sometimes even contradictory. Therefore, the aim of the review was to systematize and analyze the data that influence the in vivo magnetofection processes after the systemic injection of magnetic nanostructures. The main emphasis is placed on the structure and coating of the nanomagnetic vectors. The present problems and future trends of the method development are also considered.

**Keywords:** magnetofection in vivo; magnetic nanoparticles; iron oxide; gene delivery; gene vectors

#### **1. Introduction**

The development of nanosystems that effectively deliver genes into a cell in vivo using magnetofection is a complex and urgent task. A solution to this problem will lead to a significant progress in the creation of drug formulations for gene therapy. In such kind of therapy, nucleic acid molecules delivered into a cell can be used for over-expression of a desired protein, for gene knock down effects, for bypassing or even reparation of genetic mutations, or for activation of the innate immune system [1,2]. There are two main factors limiting biomedical applications of the technology. The first one is not fully understood the process of nucleic acids uptake, their intracellular interactions, intracellular trafficking and regulation of nucleic acid action inside cells at the molecular level. The second one is the direct delivery of nucleic acids to target cells [1,3]. Therefore, choice of the reliable and efficient delivery vectors is very important. When we talk about nonviral vectors, magnetic nanoparticles (MNPs) are often meant. They are considered to be in the center of nanotechnology-based structures for nanomedicine [4,5]. Due to their ability to respond to the magnetic field, MNPs have become very attractive for different theranostic applications [6–8]. When using MNPs, it becomes possible to carry out the magnetically controlled accumulation and release of these particles [9–11], execute the particle tracking by magnetic resonance imaging (MRI) [12–14], imaging of tumors [15,16], precise and quantitative monitoring in vitro [17–19] and in vivo [20,21], by the magnetic particle quantification (MPQ) technique in an extraordinarily wide linear dynamic range (up 7 orders of magnitude). MNPs themselves can be used for tumor therapy [22,23], for targeted drug delivery to a selected part of the body [22–24], or for the magnetic separation of cells [22], as well as for magnetofection [1,25,26]. Magnetofection is defined as method for nucleic acid delivery under the influence of a magnetic field acting on nucleic acid

**Citation:** Sizikov, A.A.; Nikitin, P.I.; Nikitin, M.P. Magnetofection In Vivo by Nanomagnetic Carriers Systemically Administered into the Bloodstream. *Pharmaceutics* **2021**, *13*, 1927. https://doi.org/10.3390/ pharmaceutics13111927

Academic Editors: Marina Santiago Franco and Yu Seok Youn

Received: 18 October 2021 Accepted: 10 November 2021 Published: 14 November 2021

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vectors that are associated with magnetic nanoparticles (see scheme in Figure 1) [27,28]. After reaching the cell surface, magnetic nanosystems (usually with polyethylenimine and DNA) are internalized into intracellular vesicles called endosomes as a result of endocytosis. Moreover, for the functional delivery of nucleic acids, so-called endosome escape is required. Otherwise, the magnetic nanoconstructions will be destroyed by the cellular enzyme system. It is believed that PEI–DNA complexes escape from the endosomes due to the so-called proton sponge effect [29]. Magnetic lipoplexes behave in a similar manner [30]. As in the case of nonmagnetic PEI polyplexes [31], a three-step behavior is observed. At the first stage, the magnetic lipoplexes attach to the cell surface and slowly penetrate the cell membrane. Most lipoplexes are internalized through endocytosis during this phase. The second stage is characterized by abnormal and limited diffusion within cells. The third stage is active transport along microtubules inside the cell.

Compared to the commonly used lipid or polymer-based transfection vectors, magnetofection has several advantages, such as higher efficiency, shorter delivery time, a possibility of very local delivery [32–34]. All the above-mentioned advantages of the method are especially relevant when performing magnetofection in vivo. There are several reviews in the literature devoted to various aspects related to in vitro [1,5,27,35–41] and in vivo [42] magnetofection. The latter work considers in vivo delivery of genes using magnetic carriers injected locally to targeted tumors or tissues [42]. In the current work, we provide an overview of the published results on the systemic injection of complex magnetic nanostructures into living organisms. The delivery of drugs/genetic information to the close vicinity of the site of action (tumor, target organ, etc.) is considered to be local. The administration into the circulatory system so that the entire body is affected is regarded as systemic. The main focus is on the magnetic vector structure. We discuss in detail different types of coating of various magnetic particles, give several examples of "unusual" magnetic carriers. Special attention is given to possible ways of development of the existing technology, challenges and prospects of the method.

**Figure 1.** Magnetofection using magnetic nanoparticles. The plasmid DNA is associated with magnetic nanoparticles, which are directed, attracted, and concentrated on the surface of cell membranes, where the endocytosis process brings the nanoparticles into the cell. Adapted with permission from [43].

#### **2. Applications of Nanoscale Carriers for Magnetofection In Vivo**

This review is a logical continuation of our previous work [42], where examples of in vivo magnetofection were described in detail for the case of administrating the magnetic vectors directly into a targeted tumor or tissue. Local injections actually allow to avoid such negative effects as undesired interactions with blood components and rapid elimination from the circulation by the reticuloendothelial system (RES). The aim of the present review is to collect and analyze data that influence the processes of magnetofection in vivo after systemic injections of the nucleic acid-bearing magnetic nanoparticles. We have summarized the related results available in various publications on the topic in Table 1.

**Table 1.** Typical targets, nucleic acid types, magnetic nanoparticle compositions used in published research on magnetofection in different animals after systemic injections of magnetic carriers.

a pDNA = plasmid DNA, Antisense ODN = antisense oligonucleotide, siRNA = small interfering RNA, b combiMAG = commercial magnetofection reagent, Lp2000 = Lipofectamine®2000—commercial transfection reagent, PEI = polyethylenimine, MNB = magnetic nanobeads (MiltenyiBiotec, Auburn, CA, USA), LipoMag=commercial transfection kit, PolyMag (Chemicell, Berlin, Germany) = commercial iron oxide nanoparticles, MCL = magnetic cationic liposome, PAAIO = poly(acrylic acid)—bound superparamagnetic iron oxide, CP = polyethylenimine copolymer, Tf = transferrin, PEG = polyethylene glycol, PLL = poly-L-lysine, MNP = magnetic nanoparticle, M-MSNs = magnetic mesoporous silica nanoparticles, Gal = galactose, CMCS = carboxymethyl chitosan, NPs = nanoparticles, SPIO = superparamagnetic iron oxide, TSMCL = magnetic 1,2-dipalmitoyl-sn-glycero-3-phosphocholine (DPPC), 3b-[*N*-(*N*',*N*'-dimethylaminoethane)-carbamoyl]cholesterol (DC-Cholesterol), dimethyldioctadecylammonium bromide (DOAB) and cholesterol liposomes at a molar ratio of 80:5:5:10, MMB = micromagnetobubbles.

> The efficiency of the in vivo application of one more magnetic nanocarriers critically depends on the charge and type of the magnetic core coating. Therefore, in this section, we have divided the data into three subsections: magnetic polyplexes (magnetic nanoparticles coated with cationic polymers, Figure 2a), magnetic liposomes (lipid-coated magnetic nanoparticles, Figure 2b), and magnetic nanosystems with an additional active targeting modality (along with the response to the magnetic field, Figure 2c). Furthermore, possible ways of the method evolution are discussed in detail.

#### *2.1. Magnetic Nanoparticles Coated with Cationic Polymers*

After the development of a method to synthesize a low aggregated magnetic polyethyleneimine/DNA nanostructures (MPD) and testing these particles when injected locally into a tumor [60], the same authors continued their research on the efficient transfection activity in a serum-containing medium and MPD application in vivo after intravenous injections [47]. The magnetic vector synthesis is described in [60]. Shortly, MNPs were synthesized by two different methods in aqueous [61] and organic media [62] with subsequent modification by ligand exchange or by silane-coupling agents to prepare a negatively charged coating. For this purpose, *N*-(trimethoxysilylpropyl) ethylenediaminetriacetic acid was used. As a result, the luciferase gene expression in tumors substantially differed for the traditional transfection and magnetofection groups. The luciferase activity effected by the transfection with MPD nanostructures subjected to a magnetic field was about five times stronger compared to that obtained without field application or by the control standard

PEI-DNA complexes. It is also worth noting that a large portion of the particles in all cases settled in the animal's lungs.

**Figure 2.** Schematic examples of different magnetic nanocarriers for magnetofection in vivo: (**a**) magnetic polyplex, (**b**) magnetic liposome, (**c**) magnetic polyplex modified with transferrin. Created with BioRender.com (accessed on 7 November 2021).

In contrast, an even higher luciferase activity was recorded, when commercial magnetic nanoparticles were used, namely MNBs (MiltenyiBiotec, Auburn, CA, USA) with an average size of 200 nm [48]. The magnetic polyplex was prepared as follows: biotinylated PEI/DNA complexes were conjugated to MNBs via a sulfo-NHS LC-biotin linker. Two hours after systemic administration of the labeled MNB/Oregon green 488- PEI/pREP4/Luc nanostructures, their accumulation was assessed in vivo by a bioimaging system that detected fluorescence. When using a magnet, the intensity was about 18 times higher in the left side of the animal's chest than without the magnetic field (Figure 3), indicating that the magnet effectively attracted the nanostructures to the left chest (heart and left lung).

Additionally, the magnetic field application caused strong therapeutic gene expression in the left lung and heart. Unfortunately, there was no quantitative assessment as in another publication with similar particles [57], where the level of luciferase gene expression in the heart increased more than three orders of magnitude (!) compared with the control mice without implantation of an epicardial magnet. Another example of work using magnetic vectors based on commercial nanoparticles is discussed in [49]. To investigate whether magnetofection could be a feasible strategy for directing antisense ODN to one specific vascular site after an intra-arterial injection in vivo, the authors infused mice with Cy3-labeled antisense ODN complexed to PolyMAG magnetic particles through a femoral catheter. As a result, solely all large and small arterioles of the cremaster muscle exposed to the magnetic field after the injection of the ODN-MP mixture showed a high level of fluorescence. That was not the case in the vessels of the cremaster muscle of the contralateral testis in the same animals. In these control vessels, similar to the control animal, which was injected with a magnetic vector without applying a magnetic field, only a few large arterioles and none of the small arterioles showed fluorescence, which qualitatively indicated that magnetofection is a working method for active targeting.

**Figure 3.** Non-invasive in vivo trafficking of MNB/PEI/DNA nanostructures 2 h after systemic administration. The fluorescent image (pseudocolor) was overlaid on the photographic image. The intensity of fluorescent signal from MNB/Oregon green 488 labeled-PEI/DNA was stronger in the left chest in group with magnet (**A**) compared to the group without magnet (**B**). Adapted with permission from [48].

The work of [53] aimed at estimating the targeting capacity in vivo under the influence of a magnetic field of angiopoietin-2 small-interfering RNA (Ang-2 siRNA) plasmid/chitosancoated magnetic nanoparticles in a model of malignant melanoma (MM) in nude mice. The method of obtaining chitosan magnetic nanoparticles was quite interesting. At the first stage, magnetic Fe3O4 nanoparticles were dispersed in a chitosan solution (Zhejiang Hisun Chemical Co., Ltd., Taizhou, China) under ultrasound agitation. Subsequently, the mixture was added to a mixed phase solvent of liquid paraffin and petroleum ether supplemented with Span-80. The solution was sufficiently emulsified and agitated, then glutaraldehyde solution was slowly added dropwise. Then, the solution was incubated at 40 ◦C in a water bath for 30 min followed by adjustment pH to 9.0 with NaOH solution. The resulting solution was heated to 60 ◦C. After standing for 1 h, the precipitate was produced. Then, successive thorough washings with anhydrous ether, acetone, anhydrous ethanol, and distilled water produced the chitosan magnetic nanoparticles. As a result, it was shown, at a qualitative level, that the targeting group (particles + magnetic field) exhibited aggregation of numerous particles on the capsule of tumor tissues and inside blood vessels and staining with Prussian blue was strictly positive.

Another example of using custom-made magnetic particles is given in [51]. It was proposed to use chondroitin sulfate-polyethylenimine copolymer (CP)-coated superparamag-

netic iron oxide nanoparticles as an efficient magneto-gene carrier for microRNA-encoding plasmid DNA delivery. The magnetofection gene carrier was prepared by complexation through electrostatic interactions between CP and self-made [63] poly(acrylic acid)-bound iron oxide nanoparticles (SPIONs) (named CPIO). To evaluate the in vivo magneto-induced uptake of the magnetoplex, a biodistribution of CPIO/Cy5-DNA was studied in nude mice with U87-xenografted tumors on the right and left hind leg regions. As a result, the enhanced permeability and retention (EPR) effect [64] played a greater role in biodistribution than the magnetic field (after 48 h, both tumors glowed, the one above, where the magnet was located, glowed slightly stronger).

Magnetic mesoporous silica nanoparticles (M-MSN) are an unusual example of a magnetic vector for in vivo magnetofection. The research of [54] aimed at creating M-MSNs that differed in shape followed by a comparative study of their efficiency in suicide gene therapy of hepatocellular carcinoma (HCC). As a model, the thymidine kinase/ganciclovir of herpes simplex virus (HSV-TK/GCV) gene therapy system was used. Carboxyl-functionalized M-MSNs were first loaded with GCV. Then, PEG-g-PLL was introduced to ensure a positive surface charge of the resulting nanostructure for electrostatic absorption of the TK plasmid. At the first stage, nanosized magnetite coated with polyacrylic acid was obtained [65]. Then, spherical (S-M-MSNs) and rod-shaped (R-M-MSNs) mesoporous silica nanoparticles were synthesized by a simple sol-gel method using the obtained Fe3O4 NPs and cetyltrimethylammonium bromide (CTAB) as a template [66,67]. After that, carboxylatemodified M-MSNs were formed via reactions of M-MSNs with ammonium persulfate (APS) and further with succinic anhydride [68,69]. Finally, PEG-g-PLL and M-MSNs–COOH were covalently conjugated using a modified EDC/NHS reaction according to [70,71]. To evaluate the therapy results, the authors proposed a comparison of the sizes of the respective tumors. Compared to the control group, which was given only saline, tumor growth inhibition was observed in all other cases. In addition, the group that was treated with the magnetic field showed a significant reduction in the relative tumor volume and weight compared to the non-magnetic group. Moreover, if both types of the field (constant and variable) were used, then the tumor was less than under a single field type, while the latter, in turn, was less than without any field.

#### *2.2. Lipid-Coated Magnetic Nanostructures*

The use of magnetic lipoplexes via magnetic field-assisted systemic delivery appeared to be a much more popular method compared to the local delivery [42]. A team of authors led by Aiqiang Dong believes that liposomal magnetofection works best when administered systemically and that the magnetic field potentiates gene transfection to concentrate magnetic lipoplexes onto target cells [44–46]. In all three papers, the magnetic nanosystems had the following composition: pGFPshIGF-1Rs/combiMAG/lp2000. In [44], it was shown that the silencing efficiency of shRNAs delivered by the liposomal magnetofection after a pGFPshIGF-1R injection reached 43.4 ± 5.7%, 56.3 ± 9.6%, and 72.2 ± 6.8% at 24, 48 and 72 h, respectively, and reached an average of 43.8 ± 5.3% by lipofection. The biological distribution and target tumor suppression after magnetofection were also studied along with the potential toxicity of the method via combiMAG-carrying plasmids expressing green fluorescent protein (GFP) and short hairpin RNAs (shRNAs) targeting IGF-1R (pGFPshIGF-1Rs) in tumor-bearing mice [45]. In that work, the accumulation and delivery of pGFPshIGF-1R into tumors using magnetic nanoparticles and a magnetic field contributed to a significant decrease in tumor growth compared to the control, the suppression rate was 36% on day 30 after treatment. The same magnetic lipoplex can be used for gene therapy after heart failure [46]. The results showed that the silencing efficiency of shRNAs delivered by liposomal magnetofection reached 72.2 ± 6.8, 80.7 ± 9.6 and 84.5 ± 5.6%, at 24, 48 and 72 h, respectively, after pGFPshIGF1R injection.

Namiki Y. et al. [10] described the development of a new nanoparticle, which consisted of an oleic acid-coated magnetic nanocrystal core and a cationic lipid shell (DOTAP (*N*-[1-(2,3-dioleoyloxy) propyl]-*N*,*N*,*N*-trimethylammonium chloride) and DOPE (dioleoylphosphatidylethanolamine) (1:1)). The lipid mixture dissolved in chloroform was mixed with chloroform-based magnetic fluid. After the addition of distilled water, the homogeneous mixture was evaporated, sonicated, and purified, finally forming LipoMag. LipoMag/siRNAEGFR#4-mU (siRNA to efficiently knock down the EGFR mRNA in tumor vessels) treatment under a magnetic field exhibited a ≈50% reduction in the tumor volume compared with the control group on day 28th after the treatment initiation.

It is known that liposomes composed of DC-Chol(3β-[*N*-(*N*',*N*'-dimethylaminoethane) carbamoyl] cholesterol) and DOPE (Dioleoylphosphatidylethanolamine) have been classified as one of the most efficient vectors for transfection of pDNA into cells and in clinical trials [72–74]. In the work [50], DC-Chol and DOPE (1:1 molar ratio) were chosen as the liposome composition, MAG-T (aqueous dispersion of magnetite Fe2O3; tartaric acid matrix; mean diameter of 20 nm) was used as the core, and the MCLs were prepared using reverse-phase evaporation. Since the target organ was the liver, the observed luciferase activity 24 h after injection was only one when using MCLs/pDNA nanoformulations under the influence of a magnetic field and 1.5-fold higher compared to the same particles without the field (Figure 4).

**Figure 4.** Effect of external magnetic field on the transfection activity of MCLs/pDNA nanoformulations in vivo. The concentration of MAG-T in the MCLs was 1.5 mg/mL and the weight ratio between MCLs and pDNA was 8.0. Luciferase activity was determined 24 h post-injection in the liver, kidney, spleen, lung, and heart, respectively. Each value represented mean ± S.D. (*n* = 6). \* *p* < 0.05, # *p* < 0.01. Adapted with permission from [50].

Another interesting example employing magnetic liposomes is presented in [58]. Here, the delivery system was based on thermosensitive cationic liposomes, which were prepared with a thermosensitive cationic formulation of 1,2-dipalmitoyl-sn-glycero-3 phosphocholine (DPPC), 3b-[*N*-(*N*',*N*'-dimethylaminoethane)-carbamoyl]cholesterol (DC-Cholesterol), dimethyldioctadecylammonium bromide (DOAB) and cholesterol at a molar ratio of 80:5:5:10 (TSCL liposomes). To prepare the magnetic liposomes (TSMCL), a magnetic fluid Fe3O4 was used as the core, which was co-encapsulated with ammonium sulfate buffer into the liposomes. The authors of the study [58] designed TSMCL-DOX-shSATB1 as a combined magnetic drug targeting and a magnetofection system to improve the efficiency of simultaneous delivery of DOX and SATB1 shRNAs. Co-delivery of DOX and the shSATB1 vector in an in vivo mouse xenograft model under the influence of a magnetic field resulted in weaker tumor growth compared to control mice.

After studying the behavior of micromagnetobubbles (MMB) through a local injection [75], the same authors continued their research trying systemic administration [59]. The endothelium of the treated dorsal skin clearly showed expression of the dsRed protein 48–72 h after treatment. That expression could not be observed in the vessel wall of mice treated with the plasmid-loaded MMB, where no external magnetic field and ultrasound had been applied, or in the equally treated mice, where only a magnetic field was applied but no ultrasound. Interestingly, although mice treated with only ultrasound but without a

magnetic field showed low rates of transfection into the vascular wall in vivo, this effect was more than 60-fold lower than that observed when a magnetic field was added to retain the MMB at the vascular wall.

#### *2.3. Magnetic Nanosystems Possessing an Additional Active Targeting Modality*

In order to enhance the penetration of small interference RNA against the pololike kinase I (siPLK1) across the blood–brain barrier to treat glioblastoma (GBM), magnetic nanoparticles (Tf-PEG-PLL/MNP@siPLK1) modified with transferrin (Tf) were prepared [52] and two types of active targeting were applied. Transferrins are iron-binding blood plasma glycoproteins that control the level of free iron (Fe) in biological fluids [76]. MNPs (Fe3O4) were prepared by alkaline co-precipitation [77] with subsequent linking with Tf-PEG-PLL or PEG-PLL. When evaluating in vivo anti-GBM activity, it was found that the tumor inhibition rate raised with increasing the dosage of the magnetic nanocarrier with trans-peptide (no-field conditions were not considered). The biodistributions of another self-luminous siRNA with and without application of the magnetic field were also evaluated (Figure 5).

The fluorescence intensity was significantly higher in the brain upon administration of Tf-PEG-PLL/MNP@Cy5-siPLK1 compared to PEG-PLL/MNP@Cy5-siPLK1 without transferrin. Moreover, it turned out that the magnetic field significantly increased the accumulation of Cy5-siPLK1 in the brain tissues. That, in turn, might contribute to the high activity of Tf-PEG-PLL/MNP @ siPLK1 in vivo against GBM.

The authors of Ref. [55] developed a nanovector with double targeting properties for the efficient delivery of a tumor suppressor gene RASSF1A specifically into hepatocellular carcinoma (HCC) cells by preparing galactosylated-carboxymethyl chitosanmagnetic iron oxide nanoparticles (Gal-CMCS-Fe3O4-NPs). It is known that galactose (Gal)-modified magnetic nanoparticles can be discerned specifically by the asialoglycoprotein receptor expressed on the surface of HCC cells and specifically recognized by HCC cells [78]. Fe3O4 NPs themselves were obtained in the aforementioned research by alkaline co-precipitation [77], then coated with chitosan [79] followed by the addition of lactose and sodium cyanoborohydride to obtain Gal-CMCS-Fe3O4 NPs. The authors studied the biodistribution of the particles loaded with pcDNA6.2mir-EGFP using fluorescent microscopy to visualize GFP expression. They noted green fluorescence in liver and tumor tissues. The average efficiency of pcDNA6.2mir-EGFP transfection in liver tissue was 32.6%. In addition, the average transfection efficiency in tumor tissue was approximately 40.8% and 29.7% when using an external magnetic field and without it, respectively. No overt fluorescence was observed in sections of kidney, spleen, heart, or lung tissues. As for delivery of the RASSF1A gene for HCC treatment, the best results were achieved when a combination of the magnetic vector, external magnetic field, and intra-abdominal administration of mitomycin (MMC—chemotherapy drug) were used for treatment.

The authors of [56] considered magnetic vectors, in which the core consisted of iron oxide modified by galactose (Gal) and polyethylenimine (PEI). The latter acted as shells providing targeted delivery of therapeutic siRNA to the liver cancer. Carboxylate-capped Fe3O4 was initially synthesized via the modified oxidative co-precipitation method [80]. Subsequently, PEI was further attached to the surface of Fe3O4–COOH via 1-ethyl-3-(3 dimethylaminopropyl)carbodiimide (EDC) followed by the addition of Gal-PEG-NH2. The biodistribution study of the Gal-PEI-SPIO/Cy5-siRNA particles showed rapid accumulation of Cy5-siRNA in the liver and tumor within 8 h. The fluorescence was much more pronounced compared to the case of nano-encapsulated Cy5-siRNA. Fluorescence from Cy5-siRNA was observed within 24 h and decayed significantly over time. With Gal-PEI-SPIO nanoparticles coated with Cy5-miRNA, the fluorescence intensity began gradually decreasing only after 8 h since administration. After 24 h, it was still observable and stronger than that in the control group (RNA only). It was found that the use of a magnetic field did not play a significant role in the inhibition of tumor growth. In general, tumors treated with Gal-PEI-SPIO/si-c-Met showed an obvious decrease in volume. c-Met

is a hepatocyte growth factor (HGF) receptor that plays an important role in the proliferation, motility, differentiation, and angiogenesis of HCC [81]. As a result, both groups (with and without magnetic field), in which treatment was used with particles with siRNA, which silenced the expression of c-Met, demonstrated a significant (and almost the same) reduction in tumor volume compared to the control group.

**Figure 5.** (**A**) Biodistribution of Cy5-siPLK1 in mice organs at 12 h after the administration of Tf-PEG-PLL/MNP@siPLK1 and PEG-PLL/MNP@siPLK1 by tail vein injection. MF presents the application of magnetic field. (**B**) The statistic results of panel A. \*\* *p* < 0.01, vs PEG-PLL/MNP@siPLK1 without MF; ## *p* < 0.01, vs PEG-PLL/MNP@siPLK1 with MF. Data are expressed as the mean ± SD (*n* = 3). Adapted with permission from [52].

#### **3. Conclusions and Outlook**

In this review, we have summarized the results of publications devoted to in vivo magnetofection for the case of systemic injection of magnetic nanocarriers and tried to figure out how it works depending on the structure of these magnetic vectors. Judging from the above-cited articles, we can say that the method proved to be a good tool for delivering genetic material (pDNA, siRNA, or antisense ODN, see Table 1) to tumors and tissues via magnetic nanoparticles with various types of coating. Due to the use of a magnetic field, a rather high delivery targeting is achieved at a low loading dose. All this moves magnetofection one step closer to practical applications in gene therapy.

Indeed, it has mentioned in a number of articles that the use of a magnetic field leads to higher gene expression on qualitative [45,49,51] and quantitative [47,48,50,52,56,57] levels. There is also numerical data on a significant decrease in tumor growth compared to control groups [10,45,52,55,56,58], as well as data on silencing efficiency [44,46], which indicate the positive effect of the magnetic field on the transfection process. A logical step would be to compare the results that scientific groups, which study in vivo magnetofection, obtained with the systemic and local routes of injection of magnetic nanostructures based on the same magnetic carriers. This is what the authors of [47,60] carried out. They declared in [47] that transfection by nucleic acid-bearing magnetic nanosystems under a magnetic field produced the highest level of luciferase activity, which was approximately 5-fold higher than that of the magnetic systems without field (when administered locally, the difference was 15-fold [60]). There is also a similar example from a group that has studied the transfection properties of micromagnetobubbles in vivo (local [75] and systemic [59] injections, respectively). Unfortunately, the results of these works could not be directly compared. There are no more examples of such comparisons, probably due to the problems that may be encountered while studying magnetofection in vivo.

The first problem of the precise magnetic guidance of complex magnetic nanostructures is the effect of applied magnetic fields. It is clear that an external magnetic field of higher intensity and gradients contributes to the faster accumulation of the magnetic nanosystems in a target region. On the other hand, an excessively strong magnetic force will result in too intense aggregation of nanoparticles, thus affecting their stability and causing potential cytotoxicity [43]. In turn, a nonlinear decrease in magnetic force with an increasing distance inevitably leads to a weaker response to the external magnetic field, which may become not strong enough to steer the magnetic nanoparticles in the blood flow deep in the body [82]. It would be interesting to study the transgene expression applied with an external magnetic field of stronger gradients, which could penetrate at a larger distance on an intact animal model. In addition, it is difficult, using the current approach, to eliminate completely gene expression in other organs, at least with the commonly used NdFeB magnets. This emphasizes the need for a more focused magnetic field of a higher gradient as, for example, in magnetic systems based on NdFeB micromagnets [83]. In particular, it is reported that the magnetic targeting of deeper situated tissues or of structures situated deeper within organs is weakened, and the enhanced permeability and retention effect, in this case, plays a greater role in biodistribution than the gradient magnetic field [51,84]. To overcome this problem, one of the approaches is to use field-enhancing elements as noted in [85–88], where vascular stents were made magnetic by nickel coating. If implanting these into a deeper situated vessel, these may influence the original magnetic field when magnetized by an external magnetic field and thus may generate strong field gradients deeper into the tissues. Another example of such an invasive technique is the implantation of ferromagnetic wires and exposure of the target area or the whole patient body to an external magnetic field. In this way, strong magnetic field gradients are produced locally and allow capture and concentration of circulating magnetic nanoparticles [89,90]. Alternatively, high gradient [91] or oscillating [92–94] magnetic fields can be used for standard magnetic nanoparticles. Another option is an employment of magnetic materials with higher specific magnetization than iron oxide, but this approach raises biocompatibility concerns. A comparison between magnetite and cobalt nanoparticles (the latter exhibiting a higher magnetization than that of magnetite) of identical sizes and coatings demonstrated similar transfection efficiencies but cobalt cytotoxicity was greater, and the nanoparticles tended to aggregate [38]. For a closer look at modern magnetic systems that allow precise magnetic guiding, we direct the reader to detailed reviews [43,95–97].

The second problem is that in order to further improve the binding of ligands to specific cells, it is necessary to introduce functional ligands such as galactose, folic acid, epithelial cell adhesion molecule and α-fetoprotein that can actively interact with the corresponding binding sites on the cell surfaces [48,55,98]. For example, it is claimed in [48] that conjugation of magnetic nanosystem with a vasculature-permeabilizing agent, such as histamine [99], VEGF [100], or serotonin [101], may further help such complex magnetic nanostructures to cross the endothelial barrier and deliver therapeutic genes to other cell types.

The third problem is that magnetic nanostructures during circulation in the bloodstream inevitably encounter difficulties associated with their interactions with blood components. For example, it is known that cationic nanoparticles attract opsonizing proteins. That causes rapid plasma clearance of the nanocarriers leading to short nanoparticle plasma half-life [102]. Therefore, it is necessary to study in detail the issues related to the factors that affect blood circulation [8,103], as well as elimination [104,105] of nanoparticles for successful applications of in vivo magnetofection, search for new hybrid delivering nanomaterials [106–109] or implementation of more efficient biochemical binding and targeting techniques [110–116]. However, we believe that the magnetofection technique can become a solid option for in vivo targeting combined with improved efficacy of gene delivery, potentially in combination with other physical techniques (electroporation, sonoporation) if appropriate magnetic fields can be generated and if an ideal coating of magnetic particles is found. Still, many questions and problems remain to be addressed before it becomes an efficient and optimized clinical tool for gene therapy.

**Author Contributions:** Writing—original draft preparation, A.A.S.; writing—review and editing, A.A.S. and P.I.N.; supervision and project administration, M.P.N. All authors have read and agreed to the published version of the manuscript.

**Funding:** The research was supported by the Ministry of Science and Higher Education of the Russian Federation, contract for the organization and development of world-class research centers No. 075-15-2020-912.

**Data Availability Statement:** Not applicable.

**Conflicts of Interest:** The authors declare no conflict of interest.

#### **References**

## *Review* **Progress in Nanocarriers Codelivery System to Enhance the Anticancer Effect of Photodynamic Therapy**

**Yu-Ling Yang †, Ke Lin † and Li Yang \***

State Key Laboratory of Biotherapy and Cancer Center/Collaborative Innovation Center for Biotherapy, West China Hospital, Sichuan University, Chengdu 610041, China<EMAIL_ADDRESS>(Y.-L.Y<EMAIL_ADDRESS>(K.L.)

**\*** Correspondence<EMAIL_ADDRESS>
† YuLing Yang and Ke Lin contributed equally to the manuscript.

**Abstract:** Photodynamic therapy (PDT) is a promising anticancer noninvasive method and has great potential for clinical applications. Unfortunately, PDT still has many limitations, such as metastatic tumor at unknown sites, inadequate light delivery and a lack of sufficient oxygen. Recent studies have demonstrated that photodynamic therapy in combination with other therapies can enhance anticancer effects. The development of new nanomaterials provides a platform for the codelivery of two or more therapeutic drugs, which is a promising cancer treatment method. The use of multifunctional nanocarriers for the codelivery of two or more drugs can improve physical and chemical properties, increase tumor site aggregation, and enhance the antitumor effect through synergistic actions, which is worthy of further study. This review focuses on the latest research progress on the synergistic enhancement of PDT by simultaneous multidrug administration using codelivery nanocarriers. We introduce the design of codelivery nanocarriers and discuss the mechanism of PDT combined with other antitumor methods. The combination of PDT and chemotherapy, gene therapy, immunotherapy, photothermal therapy, hyperthermia, radiotherapy, sonodynamic therapy and even multidrug therapy are discussed to provide a comprehensive understanding.

**Keywords:** codelivery nanocarriers; photodynamic therapy; anticancer therapies; combination therapy

#### **1. Introduction of Photodynamic Therapy and Photosensitizers**

#### *1.1. Photodynamic Therapy*

Photodynamic therapy (PDT) is a modern noninvasive antitumor technique that has great clinical application potential due to its advantages of simple operation and low systemic toxicity. PDT has been clinically approved for the treatment of some tumors, such as advanced esophageal cancer and advanced lung cancer [1]. Local or systemic photosensitizers (PSs) can be activated to produce cytotoxic reactive oxygen species (ROS) after absorbing light from an appropriate wavelength laser, which can induce tumor cell necrosis and apoptosis or cause angiotoxicity to block tumor cell nutrient supply. In addition, acute inflammatory responses and immunogenic cell death (ICD) induced by PDT have been shown to activate the body's immune system and result in the reconstruction of the tumor microenvironment [2,3].

There are two main mechanisms of the photodynamic reaction. When PSs enter cells, light at a wavelength coinciding with the PS absorption spectrum irradiates the issue, and PSs will be converted from the singlet basic energy state into the excited singlet state because of photon absorption. Part of the energy directs a PS molecule to the excited triplet state. In type I reactions, the PSs in their excited triplet state react with biomolecules and ROS can be generated from radical oxygen species formed by electron transfer reactions such as superoxide ions (O2 •−) and hydroxyl radical (OH•). In type II reactions, the energy of PSs in their excited triplet state is transferred directly to the oxygen molecule in the basic energetic state, resulting in the generation of singlet oxygen, which is highly

**Citation:** Yang, Y.-L.; Lin, K.; Yang, L. Progress in Nanocarriers Codelivery System to Enhance the Anticancer Effect of Photodynamic Therapy. *Pharmaceutics* **2021**, *13*, 1951. https://doi.org/10.3390/ pharmaceutics13111951

Academic Editors: Yvonne Perrie, Marina Santiago Franco and Yu Seok Youn

Received: 7 October 2021 Accepted: 12 November 2021 Published: 18 November 2021

**Publisher's Note:** MDPI stays neutral with regard to jurisdictional claims in published maps and institutional affiliations.

**Copyright:** © 2021 by the authors. Licensee MDPI, Basel, Switzerland. This article is an open access article distributed under the terms and conditions of the Creative Commons Attribution (CC BY) license (https:// creativecommons.org/licenses/by/ 4.0/).

reactive and cytotoxic [2,4]. ROS produced by the two reactions irreversibly damage cells and microvessels, and ROS can fight against cancer by inducing apoptotic cell death, a subroutine mode of active immunity [3,5]. Type II is considered as the most important process affecting PDT efficiency. The balance of the contribution of both reactions depends on many factors, including oxygen concentration and on the structural features of the PSs.

#### *1.2. Photosensitizers*

PDT emerged as a tumor treatment in 1907, and the early progression of it is closely related to the development of PSs [4]. The PS types are generally divided into non-porphyrins and porphyrins. The typical non-porphyrin PSs are hypericin, including phenothiazinium dyes, xanthene dyes, phthalimide derivatives, indocyanine green (ICG) etc. Methylene blue (MB), toluidine blue O, Rose Bengal (RB), ICG, as well as the near-infrared dye IR780-iodide have commonly been used in recent years [6–8]. In addition, inorganic materials like graphene oxide (GO) and gold/silver nanoparticles (AuNPs/AgNPs) are also included [9–11], the latter have good biocompatibility and tunable optical properties, and 1O2 can be generated through surface plasmon excitation for PDT of tumors [12]. Metal oxides such as TiO2 and ZnO2 NPs can be photoactivated to produce ROS through electronhole pair interactions with the surrounding 1O2. These nanomaterials generally have the advantages of low toxicity, inertness, good biocompatibility and photostability [13]. Other materials, including black phosphorus nanosheets [14], fullerene [15] and GCNS [16], can also be modified to achieve multifunctionalization.

Porphyrins PSs in the early stage were mainly mixtures obtained from natural porphyrins. For example, hematoporphyrin derivative (HPD), was first used as a PS for bladder cancer in 1970 [17,18]. Unfortunately, Since such PSs had a long half-life and would accumulate excessively in the skin, patients receiving the treatment need to avoid intense light for weeks and their lifestyle is severely affected [19,20]. The second generation are substances with a pure structure, including porphyrins, chlorins, benzoporphyrins and phthalocyanines [9,10]. The common one is 5-aminolevulinic acid (5-ALA), a prodrug of protoporphyrin IX (PpIX), also the current FDA-approved standard for brain tumor visualization. Chlorin e6 (Ce6) is photoactivated at near-infrared wavelengths (664–665 nm) [21–23], has a high efficiency of singlet oxygen generation, and can penetrate deep tissue [24]. Visudyne, the active ingredient of which is verteporfin (benzoporphyrin derivative, BPD), is the only FDA-approved NIR PS for intravenous injection with high singlet oxygen yield and a low skin phototoxicity [25]. Verteporfin has poor solubility and rapid clearance speed in vivo [26], so nanocarriers (liposomes) or cosolvents are used to ameliorate these limitations [27]. Phthalocyanines have two main UV-vis absorption wavelengths: 350 nm and 680 nm, whose high intensity is one of the main characteristics of these dyes. Due to the presence of a high Q band at 680 nm, these dyes can be activated by red light with strong tissue penetration [28]. They have low toxicity and rapid elimination properties [29]. The photophysical properties of phthalocyanines are strongly dependent on central metal ions, and complexes with diamagnetic metal ions such as Zn2+ and Al3+ help to enhance PDT, including zinc phthalocyanine (ZnPc), zinc hexadecafluorophthalocyanine (ZnF16Pc), aluminum chloride phthalocyanine (ClAlPc), etc. At present, there have been a variety of phthalocyanine-based PSs in clinical trials, such as CGP55847, Photosens and so on [30]. The third generation of PSs currently under investigation is designed to synthesize substances that have a higher affinity for tumor tissue, in the form of former drugs or those based on novel drug delivery systems (liposome, polymer, or micelle) [17–19].

The way PDT induces cell death (apoptosis, necrosis, autophagy cell death, or a combination of them) is related to the localization of PSs in subcellular components. Many reports indicate that mitochondria are important targets of PDT, and PSs targeting mitochondria has been proved to induce apoptosis more effectively, which may be related to the loss of mitochondrial membrane potential, inhibition of ADP/ATP exchange, respiratory enzymes and oxidative phosphorylation [4,20,31]. PSs targeting lysosomes could induce the release of cathepsins upon photodamage, which could cleave pro-apoptotic protein Bid and pro-

mote mitochondrial apoptosis, or cleave caspase-3 and inhibit apoptosis [31]. Necrosis is caused by high levels of cell damage and usually requires a higher dose than apoptosis. When PSs is predominantly present in the plasma membrane, this is the main form of cell death. In this type of injury, plasma membrane integrity is lost, cells dissolve, and tissue inflammation is triggered. PDT has been reported to induce autophagy, a mechanism for cell survival or adaptation in which injured cells attempt to repair or remove dysfunctional organelles to promote survival. If the initial response fails, autophagy turns into a cell death signal. PDT-induced autophagy is associated with subcellular sites damaged by ROS and can be triggered by light damage to key organelles such as endoplasmic reticulum, mitochondria and lysosome [32,33].

In addition, the fluorescence capability of PSs is widely used in clinical imaging, especially NIR fluorescence imaging technology, which can depict tumor tissue in real time, thus being used to provide highly sensitive images for surgery or to monitor drug therapy [34].

#### **2. Nanotechniques to Improve Photodynamic Therapy**

Poor water solubility is a common feature of PSs and limits their clinical application. Chemical coupling combining hydrophobic PS and hydrophilic substances could enhance the circulation time in the blood [24,35]. Physical binding is a more common approach. PSs are loaded into chitosan (CS) [23,36], cationic polymer liposome microcapsules [37–39] and other carriers, which can not only improve the physical and chemical properties but also avoid degradation in the physiological environment [40–42].

The use of targeted vectors helps reduce toxicity and allows the same therapeutic effect to be achieved with smaller doses of PS. This is mainly related to increasing the selectivity of PSs tissue distribution and improving drug bioavailability [4]. In addition to targeting specific tissues, simultaneous targeting of multiple organelles has also been reported to have a stronger therapeutic effect. This can be achieved by combining PSs with different subcellular localization or by using different nanocarriers [27,43,44].

Hypoxia at the treatment site [45,46], low ROS generation efficiency, and low light penetration are also limitations. Inorganic NPs such as AgNPs can mediate greater ROS production through Ag+ ions released from the surface and have been reported to enhance the anticancer effect of PSs such as MB and Ce6 [47–49]. GO is also commonly used to enhance ROS production, and studies have found that the use of GO loaded with MB can raise the local tumor temperature to approximately 40 ◦C, effectively preventing tumor metastasis and regeneration [50]. The use of noncovalently bound graphene oxide dihydrogen porphyrin derivatives as long-wavelength absorbers of PS (λmax of 707 nm) enables light to penetrate deeper into the tumor site, overcoming the limitations of low light penetration [51]. For some hypoxic tumors, oxygen or oxygen generators can be codelivered with the PS to enhance the PDT effect through a self-supply of oxygen. Common oxygen transport carriers include nanoscale artificial red blood cell tumors targeting human serum albumin, which can deliver oxygen-loaded hemoglobin and PSs to the tumor site [52,53]. Perfluorocarbon (PFC), an artificial blood substitute that can dissolve large amounts of oxygen, can also alleviate tumor hypoxia after cotargeted delivery with PS [54]. Another method is to use metal carriers to react with peroxide to generate oxygen at tumor sites [55,56]. For example, an activated system using linoleic acid peroxide (LAHP) and iron oxide NPs (IO NPs) can induce the apoptosis of cancer cells through tumor-specific 1O2 generation and subsequent ROS-mediated mechanisms [46]. In addition, it is feasible to transform type II oxygen-sensitive photochemical reactions into Type I photochemical reactions [57]. Other approaches to overcome hypoxia include the use of siRNA to inhibit hypoxia-inducible factors.

The upregulation of vascular endothelial growth factor (VEGF) and heat shock protein 70 (HSP-70) are the main causes of PDT tolerance. To obtain a more ideal PDT treatment effect, PS can be delivered together with gene therapy drugs or small molecule inhibitors targeting VEGF or HSP-70 [58]. In addition, NO produced by photostressed tumor cells can induce anti-PDT effect and enhance tumor aggressiveness. This can be addressed by codelivery with iNOS inhibitors. These would be described in detail in the section on codelivery treatment [59–61].

#### **3. Codelivery of PSs and Anticancer Drugs with Nanoparticles**

There are many factors that contribute to decline in photodynamic therapy, including the metastatic tumor at unknown sites, inadequate light delivery, lack of sufficient oxygen and induction of an anti-PDT effect [1]. To overcome these limitations and to improve efficiency, the researchers have combined PDT with other treatments. Codelivery is one of the most common methods to achieve combination therapy and has been widely studied. Currently, common treatment schemes that are amenable to PS codelivery can be divided into the following types: chemotherapy, gene therapy, immunotherapy, photothermal therapy (PTT), hyperthermia (HT), radiotherapy, sonodynamic therapy and multiple drugs codelivery [62–71] (Figure 1). Here, we will mainly introduce the nanocarriers used for the codelivery of PSs and antitumor drugs and discuss how the drugs enhance each other.

**Figure 1.** Combination of photodynamic therapy with various therapeutic regimens and possible synergistic antitumor mechanisms. (**A**) Chemotherapy, (**B**) Gene therapy, (**C**) Immunotherapy, (**D**) Photothermal therapy, (**E**) Hyperthemia therapy/Magnetic hyperthermia therapy, (**F**) Radiotherapy, and (**G**) Sonodynamic therapy.

#### *3.1. Chemotherapy*

Chemotherapy is a commonly used method for tumor treatment, but systemic toxicity and side effects cause patients to face great pain when receiving treatment, and the efficacy is limited by stability, targeting and multidrug resistance [72]. Many reports have proven that the application of targeted delivery vectors can improve the physical and chemical properties of chemotherapy drugs, while combination with PDT can help to improve multidrug resistance and other problems [73,74]. Mechanistically, chemotherapeutic drugs bind to the DNA of tumor cells and block DNA replication, which leads to suppressed cell division and ultimately death [75]. It has also been found that many chemotherapeutic drugs can increase intracellular ROS levels and oxidation-reduction homeostasis of cancer cells, thus enhancing the sensitivity of tumor cells to PDT [76]. On the other hand, light-independent chemotherapy can kill deep-level tumor cells that PDT cannot, while increased tumor vascular permeability induced by PDT can enhance the accumulation of nanomaterials in tumors, thus enhancing the efficacy of chemotherapy [77,78].

Cationic liposomes are widely used in the codelivery system of PSs and chemotherapeutic drugs due to their preferential accumulation in the vascular endothelium [79]. The introduction of porphyrin-phospholipid (PoP), which is a kind of PS-coupled lipid, enables photoprogrammed controlled release of cationic liposomes. ROS produced by irradiation can oxidize unsaturated lipids and accelerates the release of chemotherapeutic drugs [80,81] (Figure 2A). Doxorubicin (DOX) was loaded into PoP liposomes for intravenous injection. Tumor vascular permeability and drug accumulation were significantly increased under near-infrared light, while empty PoP liposomes without drug loading also showed antitumor effects, suggesting that conjugation with phospholipids did not affect the photodynamic effect of PS [82,83]. In the synthesis of PoP liposomes, the selection of PoP and the proportion of other raw materials are crucial to the morphology and serum stability of the liposomes. Generally, a lower content of PoP in the lipid bilayer leads to higher serum stability [84]. Moreover, the dosage needs to be considered. When DOX is loaded in excess, the bilayer of PoP liposomes becomes elliptical due to instability, while loading irinotecan (a camptothecin-derived anticancer drug) does not affect the morphology of the liposomes [62]. In addition, carrier modification can achieve multiple functions; for example, 64Cu-labeled POP liposomes can be used for imaging simultaneously with treatment and have been shown to not cause drug leakage [85]. The timing of illumination after POP liposome administration is also important, and studies have shown that the drug accumulates more in tumors at short drug–light intervals [86].

Hypoxia is the cause of drug resistance in both chemotherapy and PDT. The effect of photochemotherapy can be enhanced by codelivery of oxygen carriers [87] or oxygen generators [88] loaded with therapeutic drugs. There are also substances that can serve as both chemotherapeutic agents and oxygen donors. For example, nanoplatinum (Pt), when encapsulated in liposomes with PSs, can provide oxygen as a catalase-like nanoenzyme, while Pt ions leached separately can also exert cytotoxic effects [89]. Although the application of nanocarriers can reduce the toxicity and side effects of chemotherapy drugs, it cannot improve the antitumor tolerance caused by the upregulation of heat shock protein HSP-70 after PDT. Hailong Tian [90] combined quercetin (Qu), a chemotherapy drug with the dual effects of anticancer and heat shock protein inhibition, with IR780. IR780 modified with hydrophilic biotin and Qu were assembled into a delivery carrier in solution, which successfully enhanced the therapeutic effect of PDT by inhibiting the expression of HSP-70. PDT and some chemotherapeutic agents, such as DOX and Pt, can mediate ICD. For example, the combined action of oxaliplatin and PDT was shown to expose calreticulin (CRT) on tumor cells, successfully activating host immunity and creating a suitable microenvironment for subsequent immunotherapy combination [91]. More examples of recently published studies on the codelivery of photosensitizers and chemotherapy drugs based on nanocarriers are illustrated in Table 1.

**Figure 2.** Mechanism of chemotherapy or gene therapy synergistic with PDT to enhance the antitumor effect. (**A**) Diagrammatic illustration of the mechanism of PDT-mediated controlled release of chemotherapy drugs. ROS produced by irradiation can oxidize unsaturated lipids and accelerate the release of chemotherapeutic drugs from porphyrin-phospholipid liposomes. (**B**) Diagrammatic illustration of the mechanism of photochemical internalization (PCI)-mediated endosome escape of nucleic acid drugs. When a photosensitizer is activated by an exogenous light, ROS can be produced rapidly in a short time, destroy the membrane of the intracellular body and release therapeutic genes in the cytoplasm.

*Pharmaceutics* **2021**, *13*, 1951

*Pharmaceutics* **2021**, *13*, 1951

#### *3.2. Gene Therapy*

The rapid development of nanodelivery systems has improved the penetration ability of cells and tissues and stability of exogenous genes under physiological conditions, allowing gene therapy technology to break through previous bottlenecks and enable clinical application [102,103]. The combination of gene therapy and PDT offers a high degree of precision based on complementary base pairs, which most other therapeutic combinations lack. Exogenous genes are required to efficiently enter cells and successfully escape intracellular bodies before they can play a therapeutic role, which poses great challenges to the material properties of delivery vectors [104]. Programmable vectors have been proposed to solve this problem to some extent and are commonly released under internal stimuli such as enzymes under special pH conditions. Photochemical internalization (PCI) is an intracellular delivery technique using an exogenous light as a stimulus. After activation of PS, ROS can be rapidly produced within a short time to destroy the membrane of the intracellular body and release therapeutic agents such as nucleic acid drugs in the cytoplasm [105,106] (Figure 2B). This external light-dependent regulation is more controllable and stable than the internal response-dependent system and can achieve higher spatially controlled and targeted gene delivery [107]. Another type of photoprogrammed gene regulation uses PDT active nanomaterials as gene delivery vectors, such as PPBP, the black phosphorus (BP) nanosheets prepared with PEG and PEI modification, a black scale nanomaterial that shows PDT activity under light and then specifically degrades to release siRNA in a high ROS and acidic tumor environment to achieve targeted delivery [108].

Hypoxia influences PDT, while upregulation of VEGF and HSP-70 leads to PDT tolerance. Changing the expression of these proteins by gene therapy can help resolve these challenges to some extent and enhance the sensitivity of tumor cells to PDT. HIF1α is a hypoxia-inducible factor that plays a key role in tumor cell proliferation and angiogenesis [109]. Zheng WH [110] used anisamide-targeted lipid-calcium-phosphate (LCP) nanoparticles to achieve codelivery of protoporphyrin IX (PpIX) and HIF-1α siRNA. The results showed that HIF1α downregulation not only directly inhibited tumor cell generation but also promoted ROS production in the tumor environment, thereby enhancing PDTmediated apoptosis. Nrf2 is a key antioxidation regulator that prevents ROS accumulation and promotes angiogenesis. Deng S [111] codelivered CRISPR–Cas9 ribonucleoprotein (RNP) with Ce6, ROS generated by the latter caused release of Cas9/sgRNA into cytoplasm, resulting in Nrf2 interference and preventing tumor cells escaping from ROS-mediated killing. In terms of improving PDT-induced tolerance, Jang Y [112] prepared a DOXsiVEGF-NPS/Ce6-MBS complex coloaded with VEGF siRNA and Ce6 for the treatment of squamous cell carcinoma. Under the action of VEGF siRNA, tumor angiogenesis was significantly reduced, and the antitumor effect was improved. Similar results were reported in Cao Y [113], the MnO2 nanosheet was first surface decorated with Cu2−xS and then loaded with HSP-70 siRNA to form MnO2/Cu2−XS-HSP-70-siRNA, which mediated the heat shock response and showed superior synergistic antitumor ability. In addition, the use of therapeutic genes to regulate the expression of proteins related to the growth, development, differentiation and metabolism of tumor cells can enhance the effect of PDT or supplement the limitations of PDT. For example, gene therapy can activate the body's immune system to treat metastatic cancer [114] and inhibit epithelial-mesenchymal transition (EMT) to avoid tumor recurrence [115]. Overall, the diversity of cancer-causing genes and the complexity of pathogenesis, as well as the innate differences of individuals, give gene therapy and PDT to infinite possibilities in combination. More examples of recently published studies on the codelivery of photosensitizers and gene therapy drugs based on nanocarriers are illustrated in Table 2.

*Pharmaceutics* **2021**, *13*,1951

*Pharmaceutics* **2021**, *13*, 1951

#### *3.3. Immunotherapy*

PDT can activate the body's immune response in two ways: one is to induce an acute inflammatory response of the host and release various proinflammatory signals; the other is to trigger ICD by injured or dead tumor cells to release damage-associated molecular patterns (DAMPs) and neoantigens as danger signals [3,124]. Although the immune response mediated by PDT is not enough to kill tumor cells, it can create a reconstructed immune microenvironment for further antitumor immunotherapy [63].

Immune adjuvants are a class of immune-stimulating molecules that activate tumorspecific immune responses by interacting with Toll-like receptors (TLRs) on antigenpresenting cells (APCs) [125]. The combination of immune adjuvant and PDT has the dual ability to activate the immune system. A nanometal organic framework (nMOF) formed by direct self-assembly of metal ions and PSs is often used as a codelivery carrier and is characterized by a high loading efficiency and good biocompatibility [126]. Cai Z [127] prepared PCN-ACF-CpG@HA NPs loaded with the immune adjuvant CPG by combining PS tetrakis (4-carboxyphenyl) porphyrin (H2TCPP) with zirconium ions to target tumors with high expression of the CD44 receptor. CPG and PDT together mediated a strong antitumor immune response, with significantly higher CD8+ and CD4+ T cell infiltration at the tumor site than that in the control group. Similarly, the cationic W-TBP designed by Ni K [128] based on nMOF can directly adsorb negatively charged CPG through electrostatic action, promote its internalization and DC maturation, and enhance antigen presentation in coordination with PDT-induced CRT exposure. In addition to treating tumors in situ, this codelivery combination kills distant metastatic cancer cells. Xia Y [129] evaluated the efficacy of CPG combined with the PS verteporfin in the treatment of 4T1 metastatic breast cancer; the results showed that the activation of DC cells was significantly increased, and the tumor volume of tumor-bearing mice was smaller than that of other control groups. Xu C [130] also achieved similar results in the treatment of local and metastatic B160F10 melanoma, which was enhanced by promoting DC recruitment and cytotoxic T cell infiltration at the tumor site. Exogenous antigens can also activate immune responses. Ovalbumin (OVA), a commonly used model antigen, can be used as a supplement to immune stimulation induced by PDT. The effect of such codelivery was evaluated in the studies of Huang R [131] and Ding B [132], both of which showed synergistic immune enhancement.

Immune checkpoint therapy has made significant breakthroughs in recent years and can effectively improve the immune system's response to tumors. At present, several immune checkpoint inhibitors have been approved by the FDA for clinical treatment [133]. Tumor vascular abnormalities and the immunosuppressive tumor microenvironment (TME) caused by indoleamine 2,3-dioxygenase 1 (IDO1) seriously affect the efficacy of PDTmediated immunotherapy. Codelivery of PSs with IDO1 inhibitors is beneficial for amplifying the effects of photodynamic immunotherapy. Combinations that have been reported include ferritin and polyethylene glycol–PLGA (PEG–PLGA) coloaded with ZnF16Pc and IDO inhibitor NLG919 [134] or Ce6, tyrosinase inhibitor axitinib (AXT), IDO1 inhibitor dextro-1-methyltryptophan (1MT) and human serum albumin self-assembled NPs [135]. These combinations can improve the tumor microenvironment by normalizing tumor blood vessels, improving hypoxia levels, promoting the invasion of immune effector CD8+ T cells in tumors, and reducing the immunosuppressive properties of tumors, which represents a promising tumor treatment strategy. Cytotoxic T-lymphocyte-associated antigen 4 (CTLA-4) and programmed death-ligand 1 (PD-L1) are the most common immune checkpoints and are closely related to the immune function of T cells. Xu J [136] used the self-assembly property of Ce6 and immunoglobulin G (IgG) in the nanoscale affinity range to bind Ce6 to aPD-L1 or double bind to αPD-L1 and αCTLA-4 in the immunocheckpoint blocking treatment of glioma in situ and colon cancer. This combination therapy successfully prolonged the survival of tumor-bearing mice and produced a long-term memory response, avoiding tumor recurrence. In addition, zinc phthalocyanine and aCTLA4 have been coadded into microneedles prepared by hyaluronic acid and dextran for skin cancer delivery. This minimally invasive percutaneous drug delivery platform can also effectively induce an antitumor immune response and avoid the systemic distribution of drugs to reduce toxicity and side effects [137]. Without affecting coloaded drugs, nanodelivery systems targeting important organelles such as mitochondria [138] and the endoplasmic reticulum [139] can be designed to enhance PDT-triggered ICD, thereby enhancing immune activation (Figure 3).

**Figure 3.** Mechanism by which photodynamic therapy and immunotherapy synergistically enhance antitumor effects. When photosensitizers in tumor sites are activated, they can cause acute inflammation and induce cell apoptosis or necrosis. Dendritic cells mature when stimulated by cytokines released at the site of inflammation and provide antigens to T lymphocytes in regional lymph nodes. Activated T lymphocytes become effector T cells, which are attracted to chemokines, migrate to the tumor and kill tumor cells. Different types of immunotherapy drugs play a role in different steps of the complete antitumor immune cycle. Codelivery of a photosensitizer with an immune adjuvant or tumor antigen can synergistically enhance the activation of the host immune system and improve the immunosuppressive microenvironment. The codelivery of a photosensitizer with CTLA-4 and PD-L1 monoclonal antibodies can enhance the antitumor immunity effect of T cells. The activation of these immune cells also plays a role in preventing metastasis and recurrence.

Photoimmunotherapy is a tumor-targeting therapy using specific antibodies to tumorassociated receptors chemically coupled with PSs, which is more accurate than conventional PDT and is suitable for tumors at sensitive anatomical sites. Hasan T [140] coupled epidermal growth factor receptor (EGFR) monoclonal antibody Cetuximab with benzoporphyrin derivative for pancreatic ductal adenocarcinoma treatment. The results showed that the photoimmune nanoconjugate (PIN) had a high binding specificity and could rapidly penetrate heteromorphic organoids, providing approximately 16-fold enhancement in molecular targeted NIR photodestruction. Nevertheless, single-receptor targeted therapy may cause tumor subsets with low receptor expression to evade treatment and thus fail to completely ablate tumors. Hasan T [141] further constructs a triple receptor-targeted PIN (TR-PIN), cetuximab, holo-transferrin, and trastuzumab conferred specificity for EGFR, transferrin receptor (TfR), and human epidermal growth factor receptor 2 (HER-2). Researchers compared the binding ability of TR-PIN to tumor cells with different levels of receptor expression (EGFR, TfR or HER-2), and found that TR-PIN has the ability to recognize multiple tumor targets, effectively photodynamically eradicating different tumor subsets and reducing escape. More examples of recently published studies on the codelivery of photosensitizers and immunotherapy drugs based on nanocarriers are illustrated in Table 3.

212

**Table3.**Codeliveryofphotosensitizersandimmunotherapydrugsbasedonnanocarriers.

#### *3.4. Photothermal Therapy (PTT)*

Additionally, as a minimally invasive treatment, PTT has certain similarities with PDT. Photothermal agents concentrated at the tumor site absorb laser radiation energy and convert light energy into hyperthermia, resulting in the thermal ablation of adjacent cells. PTT enhances the effect of PDT, mainly by improving blood flow and increasing oxygen content in tumors [146] (Figure 4A).

**Figure 4.** Mechanism of photodynamic therapy combined with photothermal therapy, hyperthermia and magnetic hyperthermia in synergistic enhancement of the antitumor effect based on codelivery systems. (**A**) When light-activated photothermal agents, (**B**) or alternating magnetic fields are applied to magnetic nanoparticles, heat is generated at the tumor site, increasing blood flow and synergistically enhancing PDTSome photothermal agents of inorganic materials themselves can be used as nanocarriers, which makes the combination of PTT and PDT convenient. Gold nanorods have been used as carriers with photothermal conversion capability, and the joint delivery of DOX and PS ICG can simultaneously achieve

chemo-PTT-PDT triple therapy [69]. Gold nanocages [147] show strong absorption in the NIR region, and their empty interior and porous walls are suitable for encapsulating PSs. Bo Tian et al. [148] coupled Ce6 on PEG-functionalized GO (GO-PEG-CE6). The photothermal effect of GO promoted the transfer of Ce6, and its destruction effect on cancer cells was significantly better than that of free Ce6. The photothermal agent plasma copper sulfide (Cu2-XS) has also been reported to have photodynamic properties [149], allowing the combination of PTT and PDT to be achieved simultaneously, and has been verified in in vitro cultured melanoma cells and mouse melanoma models. The poor photostability and potential long-term toxicity of inorganic nanomaterials limit their clinical application, so more materials with good biocompatibility and high stability have been developed as alternatives. The conductive polymer polypyrrole is a material with relatively high biocompatible photostability and photothermal conversion efficiency, using polyacrylamide (PAH), polyacrylic acid (PAA) and AlPCS4. Further modification of polypyrrole resulted in a more stable AlPCS4@PPyCONH-PAH-PAA nanoneedle complex in a physiological environment. The results showed an enhanced synergistic effect, with tumor ablation in mice 14 days after treatment and no recurrence within 30 days [150].