Research ArticleHEALTH AND MEDICINE

Immune cell shuttle for precise delivery of nanotherapeutics for heart disease and cancer

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Science Advances  23 Apr 2021:
Vol. 7, no. 17, eabf2400
DOI: 10.1126/sciadv.abf2400

Abstract

The delivery of therapeutics through the circulatory system is one of the least arduous and less invasive interventions; however, this approach is hampered by low vascular density or permeability. In this study, by exploiting the ability of monocytes to actively penetrate into diseased sites, we designed aptamer-based lipid nanovectors that actively bind onto the surface of monocytes and are released upon reaching the diseased sites. Our method was thoroughly assessed through treating two of the top causes of death in the world, cardiac ischemia-reperfusion injury and pancreatic ductal adenocarcinoma with or without liver metastasis, and showed a significant increase in survival and healing with no toxicity to the liver and kidneys in either case, indicating the success and ubiquity of our platform. We believe that this system provides a new therapeutic method, which can potentially be adapted to treat a myriad of diseases that involve monocyte recruitment in their pathophysiology.

INTRODUCTION

Hypovascularity in pancreatic ductal adenocarcinoma (PDAC) (1) and reduced blood supply to the heart following ischemic myocardial injury mean that sole reliance on drug delivery through the circulatory system is ineffective under these conditions; therefore, if this method is to be used to achieve efficient delivery of drugs to target locations, then augmentation will be needed. Vascular permeability has been used as a method of passive drug delivery (2); however, studies have shown that this phenomenon occurs only transiently in the heart and the available time window is not long enough for meaningful delivery of therapeutics (3, 4). This makes low vascular permeability a bottleneck that greatly hampers drug efficacy and deliverability. Therefore, a drug delivery platform capable of leaving the circulatory system, regardless of vascular permeability, and infiltrating deep into the disease site is attractive.

Recruitment of immune cells, such as monocytes, takes place as a natural response to a change in the physiological environment. The role of monocytes varies. In the tumor microenvironment, as a cancer-related inflammatory response, they are constantly recruited and are capable of infiltrating into the tumor site (5, 6), while after myocardial injury, splenic monocytes are recruited and are capable of infiltrating into the heart to help heal the myocardium (7, 8). Inspired by this phenomenon, we designed a lipid nanoparticle (LNP)–based drug delivery platform with an active targeting scaffold that acts as a vehicle and is capable of selectively attaching onto the surface of circulating monocytes in the blood stream, moving with them, and extravasating together with them into the diseased site.

The body consists of a myriad type of cells, and targeting a specific cell type is therefore challenging. One possible way to achieve this is to use a cell-specific ligand as the targeting scaffold. As an example, several studies have reported nanoparticles carrying macrophage-specific ligands in their cargo as therapeutics. These nanoparticles were able to deliver the ligands into the macrophages resulting in their activation (9). Although this kind of ligands can potentially be used as a targeting scaffold, we chose not to use them, as we only aim to attach our nanoparticles on the monocyte surface without activating them. Furthermore, some of the ligands may not be monocyte specific and may also target endothelial cells (10, 11), resulting in unwanted off-target accumulation. Taking these into consideration, we avoided using ligands as the targeting scaffold and we opted to use aptamers instead.

Aptamers are synthetic short, single-stranded DNA or RNA oligonucleotides used as biotechnological tools and therapeutic agents. They can be designed to have high affinities toward specific proteins through their folding into tertiary structures (12). The idea of using oligonucleotides to target proteins emerged in the early 1990s, and since then, aptamers have been widely applied in many fields, including food safety, environmental monitoring, clinical diagnosis, and therapy (12). With the development of cell systematic evolution of ligands by exponential enrichment (Cell-SELEX), it has become possible to design and select aptamers with high affinities toward specific cells types, such as monocytes, while avoiding unwanted bindings to endothelial cells (13). In this study, we took advantage of this advanced technique to select a specific monocyte-targeting aptamer and integrated it with our LNP as an active-targeting scaffold to produce a high-affinity monocyte-targeting drug delivery vehicle.

Several studies have described a similar strategy whereby the body’s own cells were used to carry nanoparticles to diseased sites. T cells carrying nanoparticles loaded with a topoisomerase inhibitor ligand SN-38 were reported to reduce tumor burden in mice with disseminated lymphoma (14). LNPs carrying tumor necrosis factor–related apoptosis-inducing ligand were able to attach onto the surface of leukocytes and kill colorectal and prostate cancer cells, as well as circulating tumor cells in mice (15). Furthermore, by hitchhiking on the surface of red blood cells, nanogels carrying reteplase, a thrombolytic enzyme, ameliorated pulmonary embolism in mice (16). Our strategy, on the other hand, makes use of monocyte recruitment to the diseased site. We hypothesize that because the recruitment is an active process, it ensures that the nanoparticle and its cargo can reach the site it is intended. We also hypothesize that our monocyte-targeting drug delivery platform is versatile and can be used to treat myocardial ischemia-reperfusion (IR) injury and pancreatic cancer, two very different deadly diseases, which involve the monocyte recruitment phenomena that we harness in our strategy.

IOX2, a potent and selective hypoxia-inducible factor (HIF)–1α prolyl hydroxylase–2 inhibitor, is capable of preventing proteasome-mediated degradation of HIF-1α (17, 18). The HIF-1α protective effect of IOX2 not only contributes to the reduction of apoptosis but also enhances the transcription responses of HIF-1α (19, 20). Gemcitabine is a common chemotherapeutic agent for pancreatic cancer. It is a deoxycytidine analog capable of inhibiting the DNA replication in cancer cells and causing cell death (21). We encapsulated both of these drugs separately into our delivery vehicle, and by doing so, we were able to successfully ameliorate IR injury (using IOX2-loaded nanoparticles) and reduce tumor burden in PDAC mice (using gemcitabine-loaded nanoparticles). Moreover, unlike other bio-based materials, our aptamer-based scaffold is not patient specific, synthetic, and can be chemically modified, which are highly advantageous traits in the clinical setting.

RESULTS

Injury and tumor-induced recruitment of monocytes

As our delivery of therapeutics to disease sites relies on the recruitment of monocytes, we first examined the most efficient time point for delivery by constructing monocyte recruitment profiles to the injured heart and tumor site using IR (Fig. 1, A to C) and PDAC (Fig. 1, D to F) models of transgenic CCR2RFP/+ mice, respectively. We observed an increase in the number of recruited monocytes following IR injury and PDAC model establishments, which reached a maximum at day 4 after IR injury (Fig. 1B) and day 7 after KPC (KrasG12D, p53fl/fl, Pdx1-Cre) tumor cell transplantation (Fig. 1E). Furthermore, the number of circulating monocytes after IR injury and KPC tumor cell transplantation showed significant difference until 5 hours and day 14, respectively (figs. S1 and S2). Recruitment of monocytes to the IR heart was further confirmed by fluorescence-based intravital microscopy of the heart, whereby CCR2RFP/+ monocytes were observed (Fig. 1C). In the PDAC model, transplantation success and recruitment of monocytes were further confirmed by fluorescence-based intravital microscopy, whereby green fluorescent protein (GFP)+ KPC cells and red fluorescent protein (RFP)+ CCR2 monocytes were clearly observed at the injection site (Fig. 1F).

Fig. 1 The J10 aptamer specifically targets to the circulating monocytes.

(A) The in vivo imaging system (IVIS) revealed CCR2RFP/+ cell recruitment to the injured heart after IR. (B) IVIS quantification of the CCR2RFP/+ recruitment to the injured heart after IR. (C) Recruitment of CCR2RFP/+ cells in the injured heart after IR under an intravital microscope. (D) Representative IVIS images of CCR2RFP/+ monocyte recruitment in a mouse orthotopic pancreatic cancer (PDAC) model. The mouse KPC cells were luciferase and GFP double transgenic. (E) IVIS quantification of CCR2RFP/+ monocyte recruitment in the tumor site. (F) CCR2RFP/+ recruitment in the PDAC model under an intravital microscope. (G) Schematic illustration of the aptamer-based LNP delivery approach in the mouse cardiac IR and PDAC models via circulating monocytes. (H) Flow cytometric analysis of the specificity of J10 aptamer to monocyte cell lines RAW264.7 and J774A.1, as well as mouse endothelial cell line SVEC. The S2 aptamer was a random ordering of the J10 aptamer sequence. (I) Flow cytometry showed ex vivo targeting of Cy5-labeled J10 aptamer against mouse monocytes. (J) In vivo targeting of J10 aptamer–decorated quantum dots QD655 to circulating CCR2RFP/+ and CX3CR1GFP/+ monocytes via intravital imaging. (K) Polymerase chain reaction (PCR) analysis of J10 aptamer accumulation in the infarct area after cardiac IR. GAPDH, glyceraldehyde-3-phosphate dehydrogenase. One-way analysis of variance (ANOVA) with a Tukey adjustment was used to analyze data in (B) and (I). Two-way ANOVA with a Tukey adjustment was used for data analysis in (E) and (H). Unpaired Student’s t test was used to analyze data in (K). *P < 0.05, **P < 0.01, ***P < 0.001, and ****P < 0.0001. Scale bars, 100 μm (C and F) and 20 μm (K).

J10 aptamer is the potent monocyte-targeting aptamer

Aiming to produce a nanoplatform capable of binding to monocytes, we used nontoxic liposome-based nanoparticles coated with aptamers as a targeting scaffold, which are envisioned to be capable of infiltrating into the injured myocardium and pancreatic tumor site along with the monocyte (Fig. 1G). Aptamer candidates were chosen through the SELEX process against two monocyte/macrophage cell lines, RAW264.7 and J774A.1, for positive selection and the murine endothelial cell line, SVEC, for negative selection. Aptamers specific to both monocyte cell lines but not to SVEC were amplified through polymerase chain reaction (PCR). Following several rounds of SELEX, we identified aptamer J10 as the best candidate. The sequence of J10 was then scrambled to yield a control aptamer, S2 (fig. S3, A to E). The structures of both aptamers were predicted by Mfold software (22) (fig. S3, F and G). We then thoroughly investigated the capability of both aptamers to bind selectively to monocytes in vitro, in vivo, and ex vivo. Binding assays with Cy5-labeled aptamers confirmed that J10, but not S2, was capable of binding selectively to mouse monocyte cell lines (RAW264.7 and J774A.1) in vitro (Fig. 1H) and circulating myeloid (CD45+ CD11b+) cells ex vivo (Fig. 1I). Moreover, using intravital imaging to visualize the binding between circulating monocytes and QD655-labeled J10 (Fig. 1J and movies S1 to S4) clearly demonstrated that J10 selectively bound to monocytes. In vivo, intravenous injection of J10 and S2 aptamers revealed more J10 aptamer accumulated in the hearts with IR compared to S2 (Fig. 1K). J10 aptamer also has a higher binding affinity toward human monocyte cell lines THP-1 and U937, but not human endothelial cell line HUVEC, compared with S2 (fig. S4). All of these results supported our hypothesis that J10-labeled scaffold is capable of attaching selectively onto monocyte surface, which we then exploit to target the diseased sites.

J10-IOX2-LNP therapy improves cardiac function after IR injury

After we successfully identified J10 as the candidate for monocyte-targeting drug delivery platform, we then endeavored to use it as an active-targeting scaffold on the nanoparticles for the treatment of IR injury. LNPs were synthesized using a thiolated linker DNA that can readily conjugate to maleimide-containing DSPE-PEG (1, 2-distearoyl-Sn-glycero-3-phosphoethanolamine-N-[methoxy (polyethylene glycol)-2000]). The resulting DSPE-PEG–linker lipid was capable of hybridization with the aptamers (J10 or S2) to give the final monocyte-targeting LNP end product. Optimal aptamer density was determined through optimization of the molar ratio of linker:lipid, which was found to be 0.3%. A higher ratio, which translates to a higher density, did not result in a higher binding affinity to monocytes (fig. S5). Following self-assembly and encapsulation of the intended drugs (IOX2 or gemcitabine), aptamers could be decorated on the LNP surface through hybridization without conformational changes during the process (Fig. 2A) (23). Because of the complexity of the structure, mass spectrometry measurement was performed after each synthesis step to confirm the success of the synthesis and expected mass/charge ratio value was obtained for each step (fig. S6). Cryo–electron microscopy (cryo-EM) and high performance liquid chromatography (HPLC) analysis were performed to confirm successful encapsulation of IOX2 (Fig. 2, B and C). As expected, measurement of size and zeta potential showed that attachment of the aptamers increased the size and the negativity of the zeta potential following aptamer attachment (tables S1 and S2).

Fig. 2 The J10 aptamer–conjugated IOX2-LNPs target to the circulating monocytes.

(A) Step-wise synthesis of aptamer-conjugated LNPs encapsulated with IOX2. (B) Aptamer-IOX2-LNPs under a cryo-EM. Yellow arrowheads indicate precipitation of IOX2, and red arrows indicate the conjugated aptamers. Scale bars, 100 nm. (C) HPLC chromatogram of IOX2-LNPs. (D) In vitro binding affinity of aptamer-IOX2-LNPs to mouse monocyte cell lines J774A.1 and RAW264.7, as well as mouse endothelial cell line SVEC. mAU, intensity of absorbance (in milli-absorbance units); RT, retention time; ns, not significant. (E) IVIS imaging of aptamer-IOX2-LNPs accumulation in the injured heart. The particles were labeled with the DiD lipophilic cyanine dyes. (F) Quantitative analysis of DiD-labeled aptamer-IOX2-LNPs in the injured heart using IVIS. ROI, region of interest. (G) Accumulation of aptamer-IOX2-LNPs in the infarct area under an intravital microscope. The aptamer-LNPs were labeled with the DiD lipophilic cyanine dyes. Scale bars, 100 μm. (H) Biodistribution of aptamer-IOX2-LNPs in organs. One-way ANOVA with a Tukey adjustment was used to analyze data in (F). Two-way ANOVA with a Tukey adjustment was used to analyze the data in (D) and (H). ****P < 0.0001, **P < 0.01, and *P < 0.05.

Following the success of obtaining aptamer-LNPs, we examined the interaction between the LNPs and monocytes. Time-lapse live cell imaging taken over the course of 90 min of incubation between S2 and J10 aptamers with the monocyte cell line RAW264.7 showed that although some of nanoparticles were internalized, most of them remained on the surface, which is expected. More J10-LNPs were also observed on the surface of monocytes compared to S2, which further supports our finding that J10 is a better monocyte-targeting aptamer (fig. S7A and movies S5 and S6). We also investigated whether the attachment of aptamer-LNPs affected monocyte function. We profiled the cytokines [interleukin-1β (IL-1β), IL-6, IL-10, monocyte chemoattractant protein-1 (MCP-1), and transforming growth factor–β] of LNP-, J10-, and J10-LNP–treated RAW264.7 monocyte cell line using quantitative PCR. The results showed no changes in the levels of these cytokines, indicating that the nanoparticles did not affect the function of or cause adverse side effects to the monocytes (fig. S7B).

Having successfully encapsulated IOX2 in the J10-decorated nanoparticles, we then examined the ability of J10-IOX2-LNPs to bind to monocytes in vitro and to use monocytes to target IR hearts in vivo. Flow cytometry analysis using DiD [The far-red fluorescent dye DiD (1,1′-Dioctadecyl-3,3,3′,3′-Tetramethylindodicarbocyanine Perchlorate)]–labeled J10- and S2-LNPs revealed that the binding of J10-decorated LNPs to monocytes was more effective than S2-LNPs and nondecorated LNPs, with minimal binding to endothelial cells in vitro (Fig. 2D). For the in vivo study, in vivo imaging system (IVIS) analysis showed a significant increase in fluorescence for DiD-labeled J10-IOX2-LNPs compared to phosphate-buffered saline (PBS) (background) and DiD-labeled S2-IOX2-LNPs, indicative of successful targeting of J10-decorated nanoparticles to the injured hearts (Fig. 2, E and F). Intravital imaging further confirmed higher J10-IOX2-LNPs accumulation in the infarct area, suggesting that the nanoparticles successfully reached the intended site (Fig. 2G). Biodistribution study of IOX2-loaded S2- and J10-LNPs (Fig. 2H) showed a significant increase in IOX2 retention in the heart for J10-LNPs 4 hours after injection, indicating that our J10 aptamer drug delivery system successfully increased drug delivery to the heart. To confirm that J10-IOX2-LNPs delivered the IOX2 cargo by hitchhiking on the surface of monocytes, we depleted the circulating monocytes in IR mice using clodronate liposomes (24) and injected the nanoparticles. Complete blood count confirmed the success of monocyte depletion (fig. S8A), while quantification of IOX2 content in the heart showed significant decrease in clodronate-treated mice (fig. S8B). This result proved that our J10 drug delivery platform hitchhiked on the surface of monocytes to reach the injured heart.

The therapeutic effect of IOX2-loaded nanoparticles was then examined in a murine model of myocardial IR injury. The mice were injected with three doses of S2- and J10-IOX2-LNPs at 5 hours, 1 day, and 2 days after IR injury (Fig. 3A). These time points were optimal for therapy because injections at 5 hours or 5 days after IR injury resulted in a similar IOX2 accumulation level (fig. S9). Because IOX2 prevents the degradation of HIF-1α, which is up-regulated early after IR injury, early injection time points were chosen for the efficacy trial. Furthermore, because the enhanced permeability and retention effect diminishes after 24 hours (3), the fact that accumulation of IOX2 remained similar at 5 hours and 5 days further suggests that nanoparticle delivery was achieved by hitchhiking on the monocyte surface. This is also supported by our monocyte recruitment and circulating monocyte profiles (Fig. 1 and fig. S1), where the monocyte levels remained high within these time points. We then aimed to understand the drug release profile, by performing biodistribution studies of IOX2 in J10-IOX2-LNP–treated IR mice (fig. S10). The nanoparticle injection was performed 5 hours after IR injury, and the organs were collected at different time points (5 hours, 1 day, and 4 days) after injection. We found that accumulation of IOX2 was at the highest at day 1 after injection and decreased at day 4. This suggests that the body started to eliminate the nanoparticles and the drugs after 24 hours after administration.

Fig. 3 J10 aptamer–IOX2-LNPs improve heart functions after IR injury.

(A) Experimental design for in vivo functional evaluation of aptamer-IOX2-LNPs in the mouse cardiac IR injury model. (B) The protein levels of HIF-1α after aptamer-IOX2-LNP treatment. (C) Terminal deoxynucleotidyl transferase–mediated deoxyuridine triphosphate nick end labeling (TUNEL) assay for detection of apoptosis in the injured heart after aptamer-IOX2-LNP treatment. The apoptotic index was defined as of the percentage of TUNEL+ cells in a field examined. DAPI, 4′,6-diamidino-2-phenylindole; CTnl, cardiac troponin I. (D) Staining for α-smooth muscle actin (α-SMA) and isolectin IB4 (IB4) to examine the effects of J10-IOX2-LNPs on angiogenesis in the injured heart. WGA, wheat germ agglutinin. (E and F) Quantification of α-SMA+ (E) and IB4+ (F) vessels in the injured heart after aptamer-IOX2-LNP treatment. G) The effects of aptamer-IOX2-LNPs on cardiac fibrosis on day 21 after IR injury. (H) Quantification of cardiac fibrosis after aptamer-IOX2-LNP treatment. LV, left ventricle. (I to P) The effects of aptamer-IOX2-LNPs on the heart function 21 days after IR injury, including ejection fraction (EF) (I), fraction shortening (FS) (J), end-systolic volume (ESV) (K), end-diastolic volume (EDV) (L), dP/dt maximum (dP/dt max) (M), dP/dt minimum (dP/dt min) (N), ESPVR (end-systolic pressure-volume relationship) (O), and EDPVR (end-diastolic pressure-volume relationship) (P). (Q) The effects of aptamer-IOX2-LNPs on the survival rate of a mouse cardiac IR model. One-way ANOVA with a Tukey adjustment was used for data analysis. The Kaplan-Meier method and the log-rank (Mantel-Cox) tests were used for construction and analysis of the survival curves in (Q). *P < 0.05, **P < 0.01, ***P < 0.001, and ****P < 0.0001.

Following the injection of S2- and J10-IOX2-LNPs, the hearts were collected for analysis. Western blot analysis showed that J10-IOX2-LNP treatment retained the HIF-1α protein level in the heart, which indicates that IOX2 successfully reached the heart and prevented the degradation of HIF-1α. This, in turn, is indicative of a cardioprotective effect (Fig. 3B). On the other hand, terminal deoxynucleotidyl transferase–mediated deoxyuridine triphosphate nick end labeling (TUNEL) assay showed reduced number of apoptotic cells, demonstrating that our treatment prevented cardiomyocyte loss (Fig. 3C). In addition, J10-IOX2-LNP treatment also augmented angiogenesis, which was shown by the increased staining of α-smooth muscle actin (α-SMA) for vessels and isolectin B4 (IB4) for capillaries (Fig. 3, D to F). Trichrome staining of three levels of the heart on day 21 after IR injury showed that the J10-IOX2-LNP group had a significant reduction in infarct size compared to the controls, demonstrating better healing of the myocardium (Fig. 3, G and H). The results thus far indicated a better cardiac performance, which we then proved through echocardiography and cardiac catheterization experiments, which revealed that the J10-IOX2-LNP group showed significant improvement in all cardiac parameters in comparison to the control groups at day 21 (Fig. 3, I to P, and fig. S11).

To ensure the safety of our platform, we examined the hepatotoxicity [aspartate aminotransferase (AST), alanine aminotransferase (ALT) and alkaline phosphatase (ALP)] and nephrotoxicity [blood urea nitrogen (BUN) and CREA] of J10-IOX2-LNPs through serum analysis, all of which fell within the level of healthy animals (fig. S12, A to E). Histology analysis of the liver and kidneys was also performed, which showed no abnormalities (fig. S12F). All of these results combined showed that in the murine model of myocardial IR injury, our nanoparticles successfully targeted the injured hearts, resulting in improved cardiac functions, reduced infarct size, augmented angiogenesis, and, overall, prolonged survival of the mice (Fig. 3Q) without causing adverse side effects to the liver, kidneys, and monocytes.

J10-gemcitabine-LNP therapy shrinks primary and metastatic tumors and improves overall survival in PDAC mice

Following the success of IR injury treatment with our J10 aptamer delivery platform, we then continued our investigation using this platform to treat PDAC mice. Gemcitabine, a drug used for pancreatic cancer treatment, was encapsulated into the nanoparticles using a passive loading method (table S3). The encapsulation success was confirmed by cryo-EM and by exploiting the presence of nuclear magnetic resonance (NMR)–active 19F nuclei in gemcitabine using 19F NMR spectroscopy (2), as well as by HPLC (Fig. 4, A to C). A cytotoxicity assay confirmed that the gemcitabine toxicity to the tumor cells was retained following encapsulation (Fig. 4D).

Fig. 4 J10-Gem-LNPs target to the tumors via circulating monocytes.

(A) The aptamer-Gem-LNPs under a cryo-EM. Yellow arrowheads indicate precipitation of gemcitabine, and red arrows indicate the conjugated aptamers. Scale bars, 100 nm. (B) Representative 19F NMR spectrum of free and liposome-encapsulated gemcitabine. ppm, parts per million. (C) Representative HPLC chromatogram of free and liposome-encapsulated gemcitabine. (D) Cytotoxicity of free and liposome-encapsulated gemcitabine to cultured mouse pancreatic cancer (KPC) cell line. IC50, median inhibitory concentration. (E) In vitro targeting specificity of aptamer-Gem-LNPs against mouse monocyte and endothelial cell lines using flow cytometry. RAW264.7 and J774A.1 are the mouse monocyte cell lines; SVEC is a mouse endothelial cell line. (F to H) In vivo binding specificity of aptamer-Gem-LNPs to (F) monocytes, (G) lymphocytes, and (H) granulocytes. (I) Accumulation of aptamer-Gem-LNPs in mouse orthotopic pancreatic tumor determined with IVIS. The aptamer-Gem-LNPs were labeled with the DiD lipophilic cyanine dyes. (J) Quantification of gemcitabine accumulation in the mouse orthotopic pancreatic cancer using 19F NMR. Two-way ANOVA with a Tukey adjustment was used for data analysis in (E), and mixed-effects analysis was used to analyze data in (J). One-way ANOVA with a Tukey adjustment was used for data analysis in (F) to (I). *P < 0.05, **P < 0.01, and ***P < 0.001.

Having successfully encapsulated gemcitabine, we then examined the ability of J10-gemcitabine-LNPs (J10-Gem-LNPs) to selectively bind to monocytes in vitro and to deliver the cargo into the tumor site in vivo. Flow cytometry analysis using DiD-labeled J10- and S2-LNPs revealed that J10-Gem-LNPs were able to bind to monocytes more efficiently than to S2-LNPs and nondecorated LNPs, with minimal binding to endothelial cells in vitro (Fig. 4E). This was also confirmed in vivo through flow cytometry, whereby a preferential binding to monocytes but not to lymphocytes or granulocytes was observed (Fig. 4, F to H, and figs. S13 to S15). IVIS analysis of excised tumor showed that after 24 hours of LNP administration, the highest accumulation of nanoparticles was found in J10 group (Fig. 4I). We then aimed to understand the release profile of gemcitabine, by quantifying the amount of gemcitabine in PDAC mice at different time points (6, 24, and 48 hours) after injection (Fig. 4J). Comparison of gemcitabine content between S2 and J10 groups showed a significant accumulation at 24 hours and a modest accumulation at 48 hours for J10 group. This suggests that the body started to eliminate the nanoparticles and the drugs after 24 hours after administration, which is in agreement with the release profile of J10-IOX2-LNPs in IR hearts. All of these findings indicate that gemcitabine-loaded nanoparticles were able to target the tumor, with J10-decorated nanoparticles having the highest efficacy. To confirm that J10-Gem-LNPs delivered the gemcitabine cargo by hitchhiking on the surface of monocytes, we repeated the circulating monocyte depletion experiment in PDAC mice using clodronate liposomes and injected the nanoparticles. Complete blood count confirmed the success of monocyte depletion (fig. S8C), while quantification of gemcitabine content in the tumor showed a significant decrease in clodronate-treated mice (fig. S8D). This result proved that our J10 drug delivery platform hitchhiked on the surface of monocytes to reach the tumor site. Last, we investigated the effects of accumulated concentration of gemcitabine on monocytes, which showed that monocyte viability was not affected, suggesting no adverse side effects (fig. S16).

The therapeutic consequence of increased accumulation of gemcitabine-loaded nanoparticles was assessed in a murine PDAC model (Fig. 5A). TUNEL assay and proliferation assay using Ki67 showed that treatment with J10-Gem-LNPs significantly increased tumor cell apoptosis and decreased tumor cell proliferation, respectively, compared to S2-Gem-LNPs (Fig. 5, B and C), indicating that the treatment successfully hampered the growth of the tumor. This was then confirmed by IVIS and functional magnetic resonance imaging (fMRI) monitoring, which showed greater tumor growth suppression in the J10 group, in agreement with the tumor weight at the day of death (Fig. 5, D to F). Furthermore, treatment of gemcitabine-loaded nanoparticles did not affect the body weight (Fig. 5G), and serum chemistry assessment for hepatotoxicity (AST, ALT, and ALP) and nephrotoxicity (BUN and CREA) showed no adverse effects in both liver and kidney functions (fig. S17), which overall indicates the safety of J10-Gem-LNPs. All of these results combined showed that in the murine model of PDAC, our nanoparticles successfully targeted the tumor site, resulting in increased tumor cell apoptosis, reduced tumor cell proliferation and growth, and, overall, prolonged survival of the mice (Fig. 5H).

Fig. 5 J10-Gem-LNPs enhance the therapeutic effects of gemcitabine in a mouse pancreatic cancer model.

(A) Experimental design for the functional evaluation of aptamer-Gem-LNPs in a mouse orthotopic pancreatic cancer model. (B) J10-Gem-LNPs caused apoptosis of pancreatic tumor cells in vivo. The apoptotic index was determined with TUNEL assay. Scale bars, 20 μm. (C) J10-Gem-LNPs reduced proliferation of pancreatic tumor cells in vivo. The proliferation index was determined by the ratio of Ki67+ cells. Scale bars, 20 μm. (D) J10-Gem-LNPs reduced pancreatic tumor size on day 29 after treatment. The pancreatic tumor sizes were determined with IVIS to detect the luciferase activity of the mouse KPC cell line. (E) The J10-Gem-LNPs reduced pancreatic tumor size under MRI. (F) Quantification of orthotopic pancreatic tumor size harvested from mice treated with PBS, gemcitabine, Gem-LNPs, S2-Gem-LNPs, and J10-Gem-LNPs. (G) The effects of aptamer-Gem-LNPs on the body weight of the mouse orthotopic pancreatic cancer model. (H) J10-Gem-LNPs improved the survival rate of the mouse orthotopic pancreatic cancer model. (I) Effects of aptamer-Gem-LNPs on liver metastatic tumor volume under MRI. (J) Effects of aptamer-Gem-LNPs on the size of liver metastatic tumor on day 32 after treatment using IVIS. (K) Effects of aptamer-Gem-LNPs on the survival rate of mouse with liver metastatic tumors. Data in (B), (C), and (I) were analyzed with unpaired Student’s t test. One-way ANOVA with a Tukey adjustment was used for data analysis in (D) to (F) and (J). The data in (G) were analyzed with the two-way ANOVA with a Tukey adjustment. The survival curves in (H) and (K) were constructed with the Kaplan-Meier method and analyzed with the log-rank (Mantel-Cox) test. *P < 0.05, **P < 0.01, and ***P < 0.001.

As one of the most common metastatic site for pancreatic cancer is the liver, we further examined the therapeutic efficacy of our nanoparticles using a murine model of pancreatic cancer with liver metastasis (25). The progression of the metastatic tumor growth on the liver was similarly suppressed in the J10 group, as shown by fMRI and IVIS measurements (Fig. 5, I and J). Ultimately, we found that the J10-Gem-LNP platform was also capable of targeting liver metastasis, resulting in increased survival of the mice (Fig. 5K), which is in agreement to the results we obtained for the IR and PDAC models.

DISCUSSION

Previously, we have developed an injectable nanogel and reloadable targeted nanoparticles to improve the treatment of ischemic diseases such as myocardial infarction and hind limb ischemia (26, 27). However, both strategies are too invasive. Methods that rely solely on the ability of the drugs or drug-loaded nanoparticles to extravasate from the circulation into diseased sites are vastly limited by the availability and permeability of the blood vessels surrounding the sites. Although the method developed in our study also relies on the circulatory system to some extent, the drug-loaded nanoparticles were able to leave the blood stream and penetrate into the diseased site. With this strategy, we were able to successfully increase the therapeutic efficacy of drugs used in treating both IR injury and PDAC, a result that otherwise could not have been achieved.

Our aptamer-based LNP targeting system can be synthesized and is not patient specific. This eliminates the necessity to freshly prepare targeting scaffolds and, in a clinical setting, enables the treatment of patients who are in need of immediate administration of therapeutics. We have shown that our aptamer is capable of selectively binding to both murine and human monocyte cell lines (Fig. 1I and fig. S4), although the binding to human monocytes is not as strong as that to murine monocytes. This is expected, because we performed the SELEX procedure using murine monocyte cell lines, taking into account the difference between human and murine monocytes; this disparity is to be expected. Our findings have shown that circulating monocytes can be used as a shuttle bus for drug delivery using the appropriate aptamer-based targeting scaffold. Aptamers that can bind selectively to human monocytes with good affinity can be developed by following our approach using human monocytes to produce human monocyte-specific aptamers and be used for translational medicine purposes.

We have shown that our aptamer-based targeting vehicle was able to treat myocardial IR injury; however, we are limited by the monocyte recruitment time point and the number of circulating monocytes, which are at their optimum 4 days after injury (Fig. 1B and fig. S1). This time point is not early enough for the delivery of early cardioprotective therapeutics, which should ideally be administered a few hours after the IR episode. Nevertheless, delivery of therapeutics that prevents the heart from suffering further damage can be successfully achieved using our delivery method.

Using the same delivery vehicle and strategy, we assessed the therapeutic efficacy of our method in the treatment of PDAC. PDAC is known to exhibit hypovascularity, which makes treatments with reliance on the circulatory system challenging and ineffective (28). Fortunately, the development of PDAC involves the recruitment of monocytes in its pathogenesis (29), which is the basis of our therapeutic strategy. Therefore, although our aptamer-based delivery method also relies on the circulatory system to reach the tumor site, the ability of the drug-loaded nanoparticles to attach to monocytes, leave the blood vessel, and penetrate through the dense stromal extracellular matrix along with the monocytes increased the efficiency of drug delivery. This was validated by the increased amount of gemcitabine that successfully reached the tumor site, reduced tumor size and weight, and prolonged survival rate. Nevertheless, clinically, it is difficult to determine how “inflammatory” the tumor is at the time of treatment and if the treatment remains effective if given when the tumors are smaller (earlier) or larger (later). More studies involving the in vivo delivery kinetics will be required to further elucidate the therapeutic time window of this drug delivery system.

Last, our drug delivery system is potentially useful for the treatment of pancreatic cancer with liver metastasis. Before the formation of metastasis, monocytes are recruited to the liver (30, 31), to support the growth and proliferation of the invading tumor cells, in the end resulting in metastasis. Our delivery system was also assessed for treating liver metastasis, and we have shown that it was also able to reduce the metastatic tumor volume and prolong the survival of the mice suffering from pancreatic cancer with liver metastasis.

Our delivery system has a lot of advantages. It can potentially be used to deliver a wide variety of therapeutics such as small interfering RNA, modified RNA, antisense oligonucleotides, and protein drugs. It can also be used as a drug delivery platform for other diseases that involve monocyte recruitment in their pathophysiology. Furthermore, it is easy to manufacture and is not patient specific, which can potentially be useful for translational purposes. The only shortcoming of our study is that we only treated the mice for a short period of time, and although we managed to improve the overall condition and survival of the mice, we did not cure them. Prolonged treatment using our delivery platform may improve the overall outcome, and therefore, future longer-term studies are warranted.

MATERIALS AND METHODS

Animals

Male 8- to 10-week-old wild-type C57BL/6 J mice, weighing approximately 25 g, were used for all experiments, unless otherwise stated. All mice were purchased from BioLASCO or National Laboratory Animal Center, Taiwan. Mice were housed in a 12-hour day/night cycle with unlimited access to food and water. Homozygous B6.129(Cg)-Ccr2tm2.1Ifc/J (CCR2RFP/RFP) and B6.129P2(Cg)-Cx3cr1tm1Litt/J (CX3CR1GFP/GFP) mice were purchased from the Jackson laboratory, USA. Heterozygous CCR2RFP/+ and CX3CR1GFP/+ mice were generated from Institute of Biomedical Sciences, Academia Sinica, Taiwan. For both intravital imaging and monocyte profiling, 6- to 8-week-old CCR2RFP/+ mice were used, while 10- to 12-week-old CX3CR1GFP/+ mice were used for intravital imaging. All mouse experiments have been approved by Academia Sinica Institutional Animal Care and Use Committee.

Myocardial IR injury model

Mice (8 to 10 weeks old) were anesthetized with Zoletil 50 (80 mg/kg; Virbac) and Rompun (3.5 mg/kg; Bayer) and given O2 via a tracheal tube on a 37°C heating pad. The heart was accessed via left thoracotomy between the third and fourth ribs. The left anterior descending coronary artery was temporarily ligated with sutures 7-0 polypropylene through polyethylene-10 tubing for 45 min. Subsequently, polyethylene-10 tubing was removed to induce myocardial IR injury. The success of the surgery was evaluated by echocardiography on the following day.

Orthotopic PDAC mouse model and its liver metastasis model

For orthotopic tumor implantation, 5 × 105 live KPC cells suspended in 20 μl of sterile PBS were administered to 6- to 8-week-old C57BL/6 J mice by intrapancreatic injection around 2 to 3 mm from the pancreas tail. For the PDAC liver metastasis model, injection of KPC cells was performed on day 10 after orthotopic implantation by injection of 5 × 105 live KPC cells suspended in 10 μl of sterile PBS into the portal vein using a Hamilton syringe.

Manufacture of monocyte-targeting LNPs

Lipid film (total mass, 35 mg) was prepared in a round-bottom flask by dissolving 1,2-distearoyl-sn-glycero-3-phosphocholine (DSPC), cholesterol, and DSPE-PEG2000 in chloroform and DSPE-PEG2000 linker and DiD in methanol (molar ratio, 45:50:0.047:0.003:0.005). Solvent was removed under reduced pressure at room temperature, and the lipid film was lyophilized overnight.

IOX2-LNPs was prepared following a previously reported method (32). Briefly, the dry film was hydrated with 1 ml of internal buffer (200 mM calcium acetate) to form multilayer vehicles (MLVs). After the thin film was completely dissolved, the size and lamilarity of MLV were reduced by 10 freeze-thaw cycles under vacuum using liquid nitrogen and a 65°C water bath. It was then sonicated using a probe sonicator in total for 2 min through a series of 2-s sonication and 10-s pause. Following this, liposome solution was extruded through a 0.1-μm polycarbonate membrane 20 times at 65°C to obtain around 100-nm small unilamilar vehicle linker–LNP. Calcium acetate was removed using Sepharose CL-4B size exclusion column to establish the liposome cross membrane gradient. Then, IOX2 was incubated with liposome in a drug to a lipid molar ratio of 0.4 at 65°C for 30 min. Unencapsulated IOX2 was removed by Sepharose CL-4B size exclusion column with PBS as the mobile phase. Linker-IOX2-LNPs were then hybridized with J10 and S2 aptamers separately through overnight incubation at 4°C (linker:aptamer, 1:2.5). Free aptamer was removed by Sepharose CL-4B size exclusion column with PBS as the mobile phase.

For fabrication of Gem-LNPs, the dry film was hydrated by 1 ml of gemcitabine in PBS solution (75 mg/ml) to form MLV linker–Gem-LNP. After the dry film was completely dissolved, the size of MLV was reduced by 10 freeze-thaw cycles under vacuum using liquid nitrogen and a 65°C water bath. Linker-Gem-LNP was sonicated using a probe sonicator in total for 2 min through a series of 2-s sonication and 10-s pause. Linker-Gem-LNP was then extruded through a 0.1-μm polycarbonate membrane 20 times at 65°C and stored overnight at 4°C. Linker-Gem-LNPs were purified using a Sepharose CL-4B size exclusion column with PBS as the mobile phase. Pure linker-Gem-LNPs were then hybridized with J10 and S2 aptamers separately through overnight incubation at 4°C (aptamers:linker, 2.5:1), followed by purification using a Sepharose CL-4B size exclusion column with PBS as the mobile phase.

Following the encapsulation, the drug concentration was measured to be 0.0625 mg per mg/ml of lipid and 0.186 mg per mg/ml of lipid for IOX2 and gemcitabine, respectively. The dosages used for the in vivo experiments are 0.7 mg of IOX2/kg for three injections and 1.66 mg of gemcitabine/kg for three injections.

Intravital imaging

The multiphoton intravital imaging was performed following a published procedure (33). All animals were anesthetized by 1.5% isoflurane (Minrad) during the experiment. Injection of 100 μl of 5 mM S2-IOX2-LNP and J10-IOX2-LNP was administered to IR day 1 CCR2RFP/+ mice for an hour, and then the infarct area was visualized by a multiphoton microscope (FVMPE-RS, Olympus). Because the fluorescence of DiD-labeled IOX2-LNP was quenched within seconds under multiphoton imaging, QD655s (20 μl; Invitrogen) modified with S2 or J10 were injected to CCR2RFP/+ and CX3CR1GFP/+ mice to visualize J10-QD655s–tagged monocytes passing through the blood vessel.

Statistics and software

GraphPad Prism 8 was used for all statistical analysis and graph generation. Statistical tests are described in the figure legends. For group analysis, one-way or two-way analysis of variance (ANOVA) with Tukey’s multiple comparison tests was used. For survival analysis, deaths were recorded and used to generate Kaplan-Meier survival curves, which were compared using Mantel-Cox log-rank tests. IVIS images of tumor luminescence and nanoparticle fluorescence were quantified using Living Image 3.1 software. For tumor size quantification, MRI images were processed in Avizo using the measure tool. 19F NMR spectra acquisition was performed on Bruker TopSpin 2.1 and processed on Bruker TopSpin 2.1 or 4.0.2. Adjustments to immunofluorescence image brightness and contrast were made to improve visual clarity and were applied equally to all images within a series. Figures were assembled in Adobe Illustrator.

SUPPLEMENTARY MATERIALS

Supplementary material for this article is available at http://advances.sciencemag.org/cgi/content/full/7/17/eabf2400/DC1

https://creativecommons.org/licenses/by-nc/4.0/

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REFERENCES AND NOTES

Acknowledgments: We would like to thank the aptamer core facility in the Institute of Biomedical Sciences (IBMS), Academia Sinica for Cell-SELEX assistance. We would also like to thank the IBMS Flow Cytometry Core facility for flow cytometry analysis and Y.-H. Chen and IBMS Animal Core staff for animal experiments. We thank Academia Sinica High-Field NMR Center (HFNMRC) for technical support. We also thank J.-H. Lin, P.-J. Lin, and S.-C. Ruan DVM for assistance with the animal experiments. Funding: This work was supported by the Ministry of Science and Technology, Taiwan (MOST 108-2319-B-001-004, 108-2321-B-001-017, and 108-3111-Y-001-053), the National Health Research Institutes grant EX109-10907SI and the Academia Sinica Program for Translational Innovation of Biopharmaceutical Development-Technology Supporting Platform Axis (AS-KPQ-106-TSPA), the Thematic Research Program (AS-107-TP-L12), and the Summit Research Program (MOST 107-0210-01-19-01). HFNMRC is funded by the Academia Sinica Core Facility and Innovative Instrument Project (AS-CFII-108-112). Author contributions: S.-S.H. and K.-J.L. designed and performed experiments and contributed to data analysis, manuscript, and figure preparation. H.-C.C. contributed to the data analysis, discussion, and figure design. R.P.P. performed experiments and contributed to the discussion and manuscript preparation. C.-H.H., O.K.C., S.-C.H., and C.Y.B. performed experiments. C.-B.J. and X.-E.Y. contributed to the IOX2-liposome fabrication. D.-Y.C. and C.W.K. performed the intravital imaging. T.-C.C. established the orthotopic pancreatic cancer model. L.-L.C. drew the schematic illustration. J.J.L. and T.J.K. contributed to the discussion. P.C. managed the intravital imaging. Y.-W.T. contributed to the discussion of PDAC experiments. H.-M.L. managed the liposome fabrication and characterization. P.C.-H.H supervised and managed the project. Competing interests: T.J.K. serves as a consultant for Fujifilm Cellular Dynamics Incorporated. P.C.-H.H., S.-S.H., K.-J.L., and H.-C.C. have patent provisional applications (US 2020/63030674 and US 2020/63030555) related to the use of aptamer-based drug delivery for treatment of heart diseases and cancer. The patent provisional applications were filed by Academia Sinica. The authors declare that they have no other competing interests. Data and materials availability: All data needed to evaluate the conclusions in the paper are present in the paper and/or the Supplementary Materials. For patent and tech transfer concerns, the raw and analyzed datasets generated during the study are available for research purposes from the corresponding author on reasonable request. Additional data related to this paper may be requested from the authors.

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