Size-Modulable Nanoprobe for High-Performance Ultrasound Imaging

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Size Modulable Nanoprobe for High-Performance Ultrasound Imaging and Drug Delivery Against Cancer Lu Zhang, Tinghui Yin, Bo Li, Rongqin Zheng, Chen Qiu, Kit S. Lam, Qi Zhang, and Xintao Shuai ACS Nano, Just Accepted Manuscript • DOI: 10.1021/acsnano.8b00076 • Publication Date (Web): 10 Apr 2018 Downloaded from http://pubs.acs.org on April 10, 2018

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Size Modulable Nanoprobe for High-Performance Ultrasound Imaging and Drug Delivery Against Cancer Lu Zhang†,‡,§,#, Tinghui Yin†,#, Bo Li§, Rongqin Zheng*,†, Chen Qiu†, Kit S. Lam*,‡, Qi Zhang†, Xintao Shuai*,†,§

†

Guangdong Provincial Key Lab of Liver Disease and Department of Medical Ultrasonic,

The Third Affiliated Hospital of Sun Yat-sen University, Guangzhou 510630, China ‡

Department of Biochemistry and Molecular Medicine, UC Davis Comprehensive Cancer

Center, University of California Davis, Sacramento, California 95817, USA §

PCFM Lab of Ministry of Education, School of Materials Science and Engineering,

Sun Yat-sen University, Guangzhou 510275, China #

These authors contributed equally to this work.

* Correspondence: [email protected], [email protected], [email protected].

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ABSTRACT: Among medical imaging modalities available in the clinic, ultrasonography is the most convenient, inexpensive, ionizing radiation-free and most common. Micron-size perfluorocarbon

bubbles

have

been

used as

efficient contrast for intravascular

ultrasonography but they are too big for tumor penetration. Nanodroplets (250-1000 nm) encapsulating both perfluorocarbon and drug have been used as ultrasound-triggered release drug delivery platform against cancer, but they are generally not useful as tumor imaging agent. The present study aims to develop a type of pH-sensitive, polymersome-based, perfluorocarbon encapsulated ultrasonographic nanoprobe, capable of maintaining at 178 nm during circulation and increasing to 437 nm at the acidic tumor microenvironment. Its small size allowed efficient tumor uptake. At the tumor site, the nanoparticle swells, resulting in lowering of vaporization threshold for perfluorocarbon, and efficient conversion of nanoprobes to echogenic nano/micro-bubbles for ultrasonic imaging, and eventual release of doxorubicin from the theranostic nanoprobe for deep tissue chemotherapy, triggered by irradiation with low frequency ultrasound.

KEYWORDS: ultrasonographic nanoprobe, theranostics, perfluorocarbon, size regulation, pH-/thermosensitivity, fluorescence/ultrasound imaging, deep tissue chemotherapy

2


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Clinical diagnosis of cancer has been advanced greatly by modern imaging techniques such as computed tomography (CT), medical magnetic resonance imaging (MRI), ultrasonography.1-5 New anticancer drugs are being developed each year, which is bringing tremendous hopes to cancer patients. Yet, there are no clinically approved efficient probes for concurrent delivery of both imaging agents and effective drugs to tumors. Such theranostic probes will allow doctors to monitor the organ distribution and therapeutic effect of administered drugs in a noninvasive manner, which may potentially lead to better therapeutic outcomes. To date, many theranostic nanoprobes are being developed but none has been clinically approved.1, 6-14 The imaging sensitivity and controlled drug release property are two critical considerations for the design of efficacious theranostic nanoprobes for image-guided drug delivery. Among medical imaging modalities available in the clinic, ultrasonography is the most convenient, inexpensive and ionizing radiation-free.15, 16 In addition, it is the most common real-time imaging modality used in the clinic. Unfortunately, ultrasonography is generally associated with low detection sensitivity, which greatly limits its applications.17 In recent years, lipid-based gas-filled microbubbles (MBs) have been developed as probes for ultrasound (US) imaging. In order to have high echogenicity for ideal constrast-enhanced imaging, MBs are micron-scale sizes (1-8 µm), e.g. SonoVue averaging about 2.5 µm.18-21 They have been widely used as blood pool agents in clinic to image blood perfusion inside organs and blood flow rate inside vessels. These images are important for clinical evaluation of cardiovascular diseases, and angiogenesis or angiostenosis in diseased tissues or organs.22-24 In principle, MBs have the potential to be used as US-sensitive delivery systems to control drug release at tumor sites. Upon US exposure, MBs can oscillate and collapse rapidly, causing rapid drug release from MBs and generation of shock waves, shear forces and 3


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microstream. Such sonoporation effect on extracellular matrix and cell membrane may enhance tissue penetration and tumor cell uptake of released drugs.25-27 Unfortunately, MBs given intravenously cannot be used as theranostic probes for solid tumors, because leaky tumor neovasculatures only allow particles smaller than 700 nm to pass through,28-30 and significantly smaller nanoparticles (< 200 nm) have better chance to accumulate at solid tumors via an enhanced permeability and retention (EPR) effect.31,

32

Several nanoscale

echo-contrast agents such as perfluorocarbons (PFC)-filled nanoemulsion or nanodroplets with sizes ranging between 250-500 nm have been reported in the last decade.33-43 The shells of these nanoscale echo-contrast agents are usually comprised of phospholipid, albumin or block copolymers, and prepared through emulsification method. Rapoport et al., reported a different preparation approach by first preparing polymeric micelles using amphiphilic block copolymer, followed by emulsification on ice under low-frequency ultrasound with addition of PFC to form nanodroplets, with size ranging between 250-750 nm. For echo-detection or ultrasound-triggered drug release, the PFC encapsulated nanodroplets need to be vaporized into gas bubbles. Vaporization threshold is higher for smaller droplets because surface tension at the interface between a droplet and the bulk liquid is higher for smaller particle, which in turn leads to higher Laplace pressure inside the droplet, where the PFC resides.44 Although tumor accumulation of nanocarriers smaller than 200 nm is more efficient, vaporization threshold for these small nanocarriers (micelles) are generally too high and impractical for clinical applications. Most of the reported work thus far has been in the area of ultrasound-triggered drug release from the nanodroplets, rather than on using these nanodroplets as both imaging and drug delivery agents.38, 43-45 The present study aims to develop a type of polyamino acid 4


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polymersome-based theranostic ultrasonographic nanoprobe combining the advantages of nanoparticles ('small' in bloodstream for tumor accumulation) and microbubbles ('big' and gas-filled in tumor for high echogenicity) (Figure 1). The size of such nanoprobes encapsulating perfluoro-n-pentane (PFP), pentafluorobutane (PFB) and doxorubicin (DOX·HCl) is less than 180 nm in the circulation (pH=7.4), and can accumulate in the tumors much more efficiently than the larger nanodroplets (250-400 nm). Once getting inside the tumor via EPR effects, the nanoprobes swell to 437 nm because of the acidic tumor microenvironment (pH≈6.8). Such size increase results in lowering of the intradroplet Laplace pressure and therefore lowering the PFC vaporization threshold. With externally applied tolerable heat, via infra-red light for surface tumor or ultrasound irradiation for deeper lesions, phase transition of encapsulated PFC occurs to form gas-filled bubbles with diameters larger than 500 nm, which are readily detectable by diagnostic ultrasound. Furthermore, continual low frequency ultrasound (LFUS) irradiation of tumor site will break the bubbles to trigger drug release. We believe such PFC filled nanoprobe can be used effectively as an on-demand ultrasound-based theranostic agent against cancer.

RESULTS AND DISCUSSION Synthesis and characterization of mPEG-PAsp(DEA-co-His-co-DIP). Diblock copolymer mPEG-PAsp(DEA-co-His-co-DIP) (PPEHD) was used as the building block for our polymersome-based nanoprobe. It was prepared by randomly grafting pH-sensitive N,N'-diethylethylenediamine (DEA), histamine (His) and 2-(diisopropylamino)ethylamine (DIP) to monomethoxy polyethylene glycol (mPEG) and poly(L-aspartic acid) (PAsp) copolymer via multistep ring-opening and aminolysis reactions (Figure 2a and Figure S1 in 5


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Supporting Information). Poly(L-aspartic acid) was chosen as copolymer backbone for the sake of nanoprobe biosafety in vivo. The desired chemical structures were verified by 1H NMR and FTIR measurements (Figure 2b, Figure S2 and Figure S3 in Supporting Information). Three different pH-sensitive groups were randomly grafted to the side chains of PAsp to endow the copolymer with desirable pH-responsiveness (Figure 2c). GPC analysis revealed a unimodal elution profile, which further confirmed the successful synthesis of copolymer (Figure S4 in Supporting Information). The copolymer, comprised of 2 kDa mPEG and 21 kDa PAsp(DEA-co-His-co-DIP) (DP≈80) (Table 1), has grafting densities of DEA, His and DIP groups at 25%, 13% and 62%, respectively. Under this compositional design, the copolymer was found to self-assemble into a nanoscale polymersome, and expands from 161±9 nm to 400±11 nm at room temperature by adjusting pH from 7.4 to 6.8. Polymersomes fabricated with conventional methods may only load hydrophobic fluorocarbons into the hydrophobic membrane,46, 47 whereas encapsulation of fluorocarbons inside the lumen is more favorable for particle expansion upon phase transition. Herein, a two-step approach was developed to load hydrophobic fluorocarbons into the inner cavity of the theranostic nanoprobe (Figure 1). In the first step, polymersome incorporating hydrophilic DOX·HCl into the aqueous core (DOX-PPEHD), with loading content of DOX·HCl at 5.11%, was prepared via the double emulsion method. In the second step, fluorocarbons were introduced into the inner core by solution exchange. In practice, the aqueous solution in the core was first replaced with isopropanol solution of PFP/PFB at pH 6.8, and then isopropanol was substituted with water at pH 7.4. The fluorocarbon mixture was used since PFP alone is insoluble in isopropanol. Most of DOX molecules could stay inside the inner core of polymersome during solution exchange because it is soluble in both water and isopropanol. 6


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The theranostic nanoprobe (PFP/PFB/DOX-PPEHD), prepared by this method, has a loading content of DOX, PFP and PFB at 3.65%, 1.96% and 1.13%, respectively. Encapsulation of PFP/PFB endowed the nanoprobe with ultrasound and thermal sensitivities, but only led to a slight size increase from 161±9 nm to 172±6 nm at pH 7.4 (Figure S5 in Supporting Information). The nanoprobe was found to be stable at 37 °C in both aqueous buffer (pH 7.4) and culture medium (pH 6.8) containing 10% fetal bovine serum (FBS), with no obvious particle size change of nanoprobes over the experimental period (Figure S6a,b in Supporting Information). Stepwise stimulations steadily regulate size and drug release behaviors of nanoprobe in vitro. Using light scattering measurement, a significant increase in particle size of nanoprobe was detected when the pH of solution was adjusted from 7.4 to 6.8 at room temperature (172±6 nm vs. 426±16 nm, Figure S5 in Supporting Information). In addition, at both pH 7.4 and 6.8, the nanoprobe only showed modest increases in particle sizes when the temperature of solution was increased from 25 °C to 37 °C (Table 2). However, remarkable increase in size was detected when the temperature of solution was increased from 25 °C to 45 °C (426±16 nm vs. 509±18 nm at pH 6.8, 172±6 nm vs. 415±10 nm at pH 7.4), owing to fluorocarbon phase transition inside the inner core of polymersome. Our results are in line with the previous finding that fluorocarbons encapsulated inside polymeric nanoparticles evaporated at solution temperatures higher than their boiling points, which are 29 °C and 40 °C for PFP and PFB, respectively.48 Above 45 °C, influence of temperatures on particle size turned weak again, due to the escape of fluorocarbon from nanoprobe. These results imply that the nanoprobe may maintain its initial nanosize while circulating in bloodstream (pH 7.4, 37 °C), which is favorable for tumor accumulation via EPR effect.49-51 Moreover, the 7


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particle size expansion of nanoprobe induced by tumor acidic microenvironment (pH≈6.8) and heating (45 °C) is favorable for achieving high echogenicity and US imaging sensitivity at tumor sites. TEM analysis further revealed the morphological and size changes of nanoprobe along with pH and temperature. When solution pH was dropped from 7.4 to 6.8, the nanoprobe swelled significantly, and increasing temperature to 45 °C caused further particle expansion without breaking the polymersomes (Figure 3a). The integrity of nanoprobe to ultrasonic perturbation at pH 6.8 was determined by TEM. After the nanoprobe solution was exposed to low frequency ultrasound (LFUS, 1 MHz) for 5 min at room temperature or 45 °C, the polymersome disintegrated into small irregular shaped nanoparticles (Figure 3a,e and Figure S7e,f in Supporting Information). The high sensitivities of nanoprobe to pH, temperature and US are crucial for its theranostical applications, i.e. US imaging and controlled drug release. We have determined that 45 °C is the optimal temperature to induce fluorocarbon phase transition to significantly enlarge the nanoprobe without breaking it. Solution pH, temperature and US irradiation were found to exert significant effect on drug release from nanoprobe (Figure 3d). Because fluorescence quenching exists for closely packed DOX molecules, the nanoprobe solution showed very weak DOX fluorescence at pH 7.4. However, at pH 6.8 and 45 °C, there was an increase in permeability of polymersome membrane leading to DOX release and a 3-fold increase in fluorescence of the nanoprobe solution. Moreover, LFUS irradiation appeared very effective in promoting DOX release. Upon LFUS irradiation, the nanoprobe solution showed at least a 5 or 6-fold increase in DOX fluorescence intensity regardless of solution temperature (45 °C or 25 °C) and pH (6.8 or 7.4) (Figure 3d and Figure S8a in Supporting Information). These results are consistent with our 8


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observation in TEM and DLS measurement that LFUS irradiation could completely destroy the nanoprobe (Figure 3e). Prolonging LFUS irradiation to 10 min or above, however, did not make any appreciable difference in fluorescence intensity of solution, suggesting that 5 min LFUS irradiation was enough to break nanoprobe for complete release of DOX (Figure S8b in Supporting Information). Nanoprobe-enhanced

tumor

imaging.

The

potential

of

fluorocarbon-loaded

polymersome as ultrasonographic nanoprobe was assessed both in vitro and in vivo. First, nanoprobe solutions were subjected to power Doppler imaging at different conditions according to TEM results (Figure 3c and Figure S10 in Supporting Information). For nanoprobe without encapsulated fluorocarbon (DOX-PPEHD), US imaging signal was either undetectable (pH 7.4 at room temperature) or very weak (pH 6.8 at room temperature or 45 °C). In contrast, PFP/PFB loaded nanoprobe (PFP/PFB/DOX-PPEHD) exhibited much stronger US imaging signal under the same condition. Changing solution pH value to 6.8 and temperature to 45 °C both significantly intensified the US imaging signal of nanoprobe solution. These data are consistent with the DLS and TEM results that showed particle size of nanoprobe was significantly increased at pH 6.8 and 45 °C, and large particles resulted from gasification possess high echogenicity. For imaging study in nude mice bearing subcutaneous xenografts of C6 glioma, a near infrared fluorescent (NIRF) dye, 1,1'-dioctadecyl-3,3,3',3'-tetramethylindotricarbocyanine iodide (DiR) replacing DOX, was incorporated into nanoprobe to enable bimodal imaging (Figure S11 in Supporting Information), and microbubble at around 1 µm was used as a conventional ultrasound contrast control. After DiR and PFP/PFB encapsulated microbubbles were injected via tail vein, DiR fluorescence reached the highest intensity in liver within 2 h 9


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after injection, which implies a fast clearance by reticuloendothelial system (RES) (Figure S12b in Supporting Information). More importantly, no tumor uptake of DiR fluorescence was detected. The ex vivo imaging of tumor and major organs from mice sacrificed at 48 h after injection further confirmed that microbubbles accumulated in liver and spleen but not in tumor (Figure S12c in Supporting Information). These results are not unexpected since microbubble is too large to cross the leaky neovasculature of solid tumors. In contrast, owing to its small size (approximately 170 nm) which is suitable for tumor accumulation via EPR effect, nanoprobe injected via tail vein effectively accumulated in tumor, regardless of fluorocarbon loading (Figure S12a in Supporting Information). The DiR fluorescent signal at tumor site reached a plateau at 11 h after nanoprobe injection, and showed no obvious decay even at 48 h, implying a persistent tumor accumulation of nanoprobe, which meant a wide time window for US imaging. The ex vivo fluorescent imaging also showed effective tumor accumulation of nanoprobe (Figure S12c in Supporting Information). The DiR fluorescence of tumor was much stronger than that of liver and other organs. Based on the in vivo fluorescence imaging results (Figure 4a), three post-injection time points 5 h, 11 h and 24 h, representing weak, strong and persistent tumor accumulation respectively, were chosen to explore the ultrasonographic potential in nude mice bearing subcutaneous C6 glioma (Figure S13 in Supporting Information). At body temperature, US imaging with power Doppler mode revealed strong echogenic signal in tumor at 11 h after tail vein injection, due to efficient tumor accumulation of nanoprobe. Consistent with the in vitro US imaging study (Figure 3c), heating tumor region to 45 °C further boosted imaging intensity, which was due to the higher echogenicity of nanoprobe at expansion driven by fluorocarbon evaporation (Figure 4b). In line with the weak tumor accumulation of nanoprobe at 5 h after tail vein injection, 10


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ultrasonography performed at this time point failed to detect a clear imaging signal at tumor site, even with heating to 45 °C. We believe our pH responsive and size-changeable nanoprobe offers great potential as contrast for ultrasonographic detection of tumor, even though it is much smaller than the micron size microbubbles. LFUS-triggered intratumor drug release and tumor-penetrating drug delivery. The LFUS-assisted drug delivery was evaluated in vivo using two nanoprobes each incorporating DiR or DOX for in vivo imaging and histological studies. To investigate LFUS-assisted drug delivery in animals bearing subcutaneous C6 glioma, the tumor was exposed to LFUS irradiation (3 times and 5 min each) at 11 h after tail vein injection of nanoprobe, at which time the nanoprobe had already reached the highest level at the tumor site (Figure 5a). At the tumor sites of animals receiving PFP/PFB/DiR-PPEHD/LFUS(+), DiR fluorescence was significantly intensified at just 5 min after LFUS irradiation (Figure 5b), indicating a fluorescence dequenching due to rapid DiR release. Retesting at 2 h after irradiation did not show much fluorescence attenuation at tumor sites. On the contrary, the DiR fluorescence intensity at tumor sites did not obviously change when nanoprobe without ultrasound sensitivity was injected or LFUS irradiation was not applied (Figure 5b,c). CLSM observation of tumor frozen sections obtained at 2 h after LFUS irradiation further revealed a tumor-penetrating drug delivery in animals receiving PFP/PFB/DOX-PPEHD/LFUS(+) treatment. The red and green fluorescence were attributed to DOX and FITC labeling tumor blood vessels. Animals receiving PFP/PFB/DOX-PPEHD/LFUS(+) treatment showed strong DOX fluorescence throughout the tumor area, whereas other animal groups only showed strong DOX fluorescence near capillaries (Figure 5d). Apparently, DOX was poorly released and hardly diffused inside tumor of animals receiving US-nonsensitive probe or US-sensitive 11


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probe but without LFUS irradiation, which was consistent with previous reports that nanoparticles poorly penetrated tumor tissue.52-54 However, in animals receiving US-sensitive nanoprobe with LFUS irradiation of tumor, DOX was released from nanoprobe and then diffused much more effectively inside tumor. In this case, the elevated US sensitivity (echogenicity) of nanoprobe due to particle expansion in acidic tumor microenvironment should have contributed to the efficient intra-tumor drug delivery assisted by LFUS irradiation. Antitumor activity of nanoprobe improved by LFUS. Finally, the therapeutic efficacy of nanoprobe was explored in nude mice bearing subcutaneous C6 glioma xenograft, and the design of our animal experiment is shown in Figure 6a. When tumors reached 50 mm3, nanoprobes were injected via tail vein, and then tumor area was exposed to local LFUS irradiation (3 times and 5 min each) at 11 h after injection. Afterwards, same administrations (i.v. injection of nanoprobes and local tumor LFUS irradiation) were repeated three times on day 4, 7 and 11. As shown in Figure 6b, the most significant antitumor effect was achieved in animals treated with PFP/PFB/DOX-PPEHD/LFUS(+). This animal group exhibited much slower tumor growth than other groups (p < 0.05). Overall, the survival curves correlate well with tumor growth results (Figure 6c). All animals receiving PBS died within 33 days, whereas 87.5% of animals receiving PFP/PFB/DOX-PPEHD/LFUS(+) treatment survived more than 40 days. The histological characteristics of tumor tissues in different animal groups were analyzed by H&E staining. Among all animal groups, the one receiving PFP/PFB/DOX-PPEHD/LFUS(+)

treatment

showed

much

lower

level

of

nuclear

polymorphism and less cancer cell density in tumor sections (Figure 6d). In addition, fluorescent staining for apoptosis obtained consistent results. Tumor sections of animals 12


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receiving PFP/PFB/DOX-PPEHD/LFUS(+) treatment exhibited much more intensive green fluorescence, indicating a high level of apoptosis. In comparison, tumor sections from other animal groups showed either no or weak green fluorescence, indicating low levels of apoptosis (Figure 6d).

CONCLUSION In summary, although lipid-based gas-filled microbubbles (MBs) have been widely applied as blood pool agents to image blood perfusion and flow rate, they are too big for tumor detection and drug delivery. Leaky vasculature of solid tumors only allows nanosized or smaller objects to pass through, but echogenicity of gas-filled nanobubbles (NBs) are too weak to enhance echocontrast in ultrasonography. A typical strategy to enhance echogenicity is to deliver fluorocarbons as smaller nanodroplets to tumor sites and then use ultrasound to induce phase transition into NBs which will coalesce transiently into larger MBs. However, the random transient nature of NB generation and coalescence makes the use of nanodroplets as ultrasound contrast agents for tumor imaging unreliable. Indeed, most work on nanodroplets relates to their use in ultrasound-triggered drug release at tumor sites rather than for tumor imaging. We believe the ultrasound-based theranostic nanoprobe reported here can achieve both. Here, we use pH-sensitive polymersome to encapsulate both liquid PFC and doxorubicin to form a 178 nm nanoprobe, which can efficiently enter tumors via EPR effects and then expands to 437 nm upon entering the acidic tumor microenvironment. Using a subcutaneous C6 glioma xenograft model, we have demonstrated that this nanoprobe has the ability to maintain nanosize suitable for long circulation time and tumor accumulation via an EPR effect.31, 54-56 Once entered the tumor where the microenvironment is acidic (pH≈6.8), this nanoprobe self-swelled, and with some local heating via infra red light 13


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or low frequency ultrasound, the encapsulated fluorocarbon underwent temperature-dependent phase-transition into gas, resulted in a significant increase in echogenicity, detectable by ultrasound image. Further irradiation of the tumor site with LFUS could trigger rapid release and tumor uptake of doxorubicin, leading to superior tumor response. Moreover, the LFUS-assisted delivery exhibited tumor-penetrating drug diffusion, which overcomes the challenge that nanocarriers may not effectively deliver drugs to tumor tissue far away from blood vessels.52,

57-59

This pH-dependent size-switchable strategy and on-demand

ultrasound-triggered theranostic nanoprobe represents a prototype of nanotheranostic agent for effective imaging and treatment of cancer with ultrasound. EXPERIMENTAL SECTION Materials. ε-Benzyloxycarbonyl L-Aspartic acid NCA, abbreviated as BLA-NCA, was synthesized

as

previously

reported.60

α-Methoxy-ε-hydroxy-poly(ethylene

glycol)

(mPEG-OH, Mn=2 kDa, Fluka) was converted into α-methoxy-ε-amino-poly(ethylene glycol) (mPEG-NH2) according to literature.61 Anhydrous N, N-dimethylformamide (DMF, ≥99.8%, Alfa Aesar), dimethylsulfoxide (DMSO, ≥99.8%, Sigma-Aldrich), dichloromethane (DCM, ≥99.7%, Sigma-Aldrich), isopropanol (≥99.5%, J&K) were used as received. N, N-diisopropylethylenediamine (DIP, ≥97%, J&K), 2-aminotriethylamine (DEA, ≥96%, Sigma-Aldrich), histamine (His, ≥95%, Sigma-Aldrich), doxorubicin hydrochloride (Zhejiang Hisun Pharmaceutical Co., Ltd., China), perfluoropentane (PFP, >98%, Sigma-Aldrich), 1,1,1,3,3-Pentafluorobutane

(PFB,

>98%,

Sigma-Aldrich),

1,1-dioctadecyl-3,3,3,3-tetramethylindotricarbocyanine iodide (DiR, AAT Bioquest) were used as received. Triethylamine (TEA) was dried using potassium hydroxide and then distilled. Dialysis tubes (MWCO: 3.5 and 14 kDa) were obtained from Shanghai Green Bird Technology Development Co., Ltd., China. PBS and fetal bovine serum (FBS) were obtained from Life Corporation (Gibco, USA). 14


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Synthesis of mPEG-PBLA. This prepolymer was synthesized by ring-opening polymerization of BLA-NCA with mPEG-NH2 as an initiator. Under argon protection, BLA-NCA (4.5 g) and mPEG-NH2 (452 mg) were dissolved in anhydrous DMF (4 mL) and dichloromethane (24 mL), respectively. After the two solutions were mixed, the polymerization was conducted at 35 °C for 72 h. After reaction, the solution was concentrated using rotary evaporation and then precipitated to cold ether. The precipitate was recovered by filtration and vacuum-dried to get mPEG-PBLA (yield: 88.6%). Preparation of mPEG-PAsp(DEA-co-His-co-DIP). 3 g of mPEG-PBLA and 378 mg of DEA were added into 15 mL of anhydrous DMSO under argon protection. The mixture was stirred at 35 °C for 24 h. Histamine (His, 181 mg) in 15 mL of anhydrous DMSO was added, and the solution was stirred at 35 °C for 24 h. After the addition of DIP (2.35 g), reaction was performed for another 8 h at the same conditions. The reaction mixture dialyzed (MWCO: 3.5 kDa,

2

h)

against

methanol

was

subjected

to

rotary

evaporation

to

get

mPEG-PAsp(DEA-co-His-co-DIP) (yield: 66.7%). Characterizations. 1H NMR spectra were recorded on a Varian Unity 300 MHz spectrometer or a Bruker 400 MHz spectrometer using DMSO-d6 or D2O to dissolve polymers. FTIR spectra of powder samples compressed to KBr pellets were recorded on a Nicolet/Nexus 670 FTIR spectrometer. The molecular weight distribution of polymer was estimated on a gel permeation chromatography (GPC) system comprising a Waters 1515 pump, an Ultrahydrogel TM 500 column, a Waters 2417 differential refractive index detector. DMF added with LiBr (1.0 g L-1) served as eluent (flow rate: 1.0 m min-1). Nanoprobe preparation. DOX-loaded nanoprobe without PFP/PFB (DOX-PPEHD) was prepared

via

double

emulsion

method.

In

brief,

after

20

mg

of

mPEG-PAsp(DEA-co-His-co-DIP) was dissolved into 2 mL of anhydrous chloroform, 200 µL aqueous solution of DOX·HCl (10 mg mL-1) was added into the polymer solution and emulsified via sonication on ice (60 Sonic Dismembrator, Fisher Scientific). The primary 15


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emulsion was further emulsified in 10 mL of PBS (pH 7.4, 0.05 mol L-1) via sonication on ice for 5 min. After chloroform was removed by rotary evaporation, the solution was dialyzed (MWCO: 14 kDa) against PBS (pH 7.4) in order to eliminate free drug. Finally, large aggregates was removed by filtering the solution through a 0.45 µm filter membrane. PFP/PFB was encapsulated into the DOX-loaded nanoprobe by solution exchange method. In brief, 1 mL of DOX-PPDHE solution (pH 6.8) was mixed with 2 mL of isopropanol at room temperature, and then 3 mL PFP/PFB (v/v = 2:1) was added to the above solution. The mixture was further shaken for 8 h at 4 °C before adding 1 mL of PBS (pH 6.8). The solution was thoroughly mixed and then kept still to allow phase separation. The water phase containing drug and fluorocarbon-loaded nanoprobe (PFP/PFB/DOX-PPEHD) was taken out with a syringe, dialyzed (MWCO: 14 kDa) against PBS (pH 7.4) to eliminate isopropanol, filtered through 0.45 µm membrane to eliminate big aggregates, and stored in refrigerator at 4 °C prior to use. Nanoprobe characterizations. Transmission electron microscopy (TEM) analysis was conducted on a Philips CM120 transmission electron microscope (Philips, Eindhoven, The Netherlands). 5 µL of sample solution (1 mg mL-1) was dried at room temperature on the copper grid with an amorphous carbon coating. A tiny drop of uranyl acetate solution (2 wt% in water) was then added to negatively stain the sample for 1 min, and then blotted off using filter paper. The grid was dried overnight before TEM analysis. The hydrodynamic sizes of nanoparticles were determined by dynamic light scattering (DLS). Measurements were preformed on a 90 Plus/BI-MAS equipment (Brookhaven Instruments Corporation, USA) with an auto-correlator detecting scattered light at a 90° angle. Three measurements for each sample were conducted. The data were presented as the mean ± standard deviation (SD). Determination of DOX loading content. After the freeze-dried PFP/PFB/DOX-PPEHD and DOX-PPEHD nanoprobes were incubated for 3 days in PBS (pH 5.0), the absorbance of DOX at 485 nm was measured with a Unico UV-2000 UV-vis spectrophotometer (Shanghai, 16


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China). The DOX loading content of nanoprobe was estimated based on a pre-established calibration curve. Determination

of

nanoprobe

multi-stimulation

sensitivity.

The

PFP/PFB/DOX-PPEHD solutions (1 mL each, nanoprobe concentration of 0.5 mg mL-1) were incubated for 2 h at the following conditions: pH 7.4, pH 6.8, pH 7.4 + 45 °C, pH 6.8 + 45 °C, pH 7.4 + LFUS, pH 6.8 + LFUS, pH 7.4 + 45 °C + LFUS, pH 6.8 + 45 °C + LFUS and pH 5.0. Room temperature was adopted for studies unless otherwise noted. The fluorescence spectrum of DOX from 500 to 750 nm for each solution was recorded at a band width of 10 nm on a PerkinElmer PE-LS55 fluorescence spectrometer (Waltham, USA). To gain the emission spectrum, a 485 nm excitation wavelength was employed. If required, nanoprobe solutions were exposed to low-frequency ultrasound (LFUS) irradiation for 5, 10 and 15 min respectively using a therapeutic ultrasound system (DCT-700, WELLD, Shenzhen, China) at 1 MHz, 50% duty cycle, acoustic pressure 2.0 MPa in the beginning of incubation. DOX release of nanoprobe in vitro. Drug release was performed at pH 7.4, pH 6.8, pH 7.4 + 45 °C, pH 6.8 + 45 °C, pH 7.4 + LFUS, pH 6.8 + LFUS, pH 6.8 + 45 °C + LFUS, pH 7.4 + 45 °C + LFUS and pH 5.0, respectively. Room temperature was adopted for release studies unless otherwise noted, and the PFP/PFB/DOX-PPEHD solution (5 mL, 1 mg mL-1) was adjusted to pH 6.8 or 5.0 with 1 N HCl. If required, the nanoprobe solution with pre-designed temperature and pH value was exposed to LFUS for 5 min before being transferred to a dialysis bag (MWCO: 14 kDa) and then dialyzed against 50 mL PBS of the same pH in an incubator shaker (ZHWY-200B, Shanghai Zhicheng, China). At pre-designed time intervals, 3 mL of solution outside the dialysis bag was replaced with fresh buffer for UV-vis analysis at room temperature (Unico UV-2000, Shanghai, China). Based on the pre-established calibration curves, DOX concentration was calculated according to the absorbance intensity of solution at 485 nm. The percentages of drug released against time were plotted. 17


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Preparation of DiR-loaded nanoprobe and microbubble. The preparation methods of PFP/PFB/DiR-PPEHD

and

DiR-PPEHD

nanoprobes

were

similar

to

that

of

PFP/PFB/DOX-PPEHD and DOX-PPEHD nanoprobes. To remove free DiR and large aggregates, nanoprobe solutions were dialyzed (MWCO: 14 kDa) against PBS (pH 7.4, 0.05 mol L-1) and filtered through a 0.45 µm filter membrane. In contrast, PFP/PFB/DiR-loaded microbubble (PFP/PFB/DiR-MB) was prepared using the thin film hydration-sonication approach. Phospholipids (18 mg DPPC, 3.5 mg PEG-DSPE and 1 mg DPPA) and DiR were codissolved in chloroform (4 mL). After natural evaporation in a 9 cm culture dish, a phospholipid thin film was formed. The film was hydrated in 4 mL of PBS at 60 °C and shaken for 1 h in an incubator at 120 rpm to form DiR-liposome. 0.5 mL of DiR-loaded liposome solution was then transferred into a 1.5 mL eppendorf tube, and the air of the container was replaced by PFP/PFB under 60 °C using a syringe equipped with a 23G fine needle. Afterwards, the eppendorf tube was sealed off and mechanically vibrated at 60 Hz using a dental amalgamator (YG-100, ZG Medical Device, Shanghai) for 45 s to allow formation of PFP/PFB/DiR-microbubble. After keeping still for 2 min, the lower PBS solution was collected into a new eppendorf tube. The eppendorf tube was centrifuged (50 × g) for 5 min. The clear solution with free DiR and free phospholipids were removed from the tube, and the remained PFP/PFB/DiR-microbubble was re-suspended in 0.5 mL of PBS. The process of centrifugation and PBS re-suspension was repeated for three times to further purify the microbubble. The final sample was stored in 4 °C refrigerator for the in vivo study. Animal models. All surgical interventions and post-operative animal care were approved by the Institutional Animal Care and Use Committee of the Sun Yat-sen University. Nude mice bearing subcutaneous xenografts of rat C6 glioma were used for in vivo studies. C6 cells were trypsinized and washed three times with PBS. Afterwards, 2.5×106 cells in 50 µL of PBS were injected into the right back of BALB/c nude mouse (4 weeks, 18-20 g). In vivo/Ex vivo fluorescence imaging. For the fluorescent imaging studies, a 18


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near-infrared fluorescent dye (DiR) rather than DOX was encapsulated into vesicle. Nude mice bearing subcutaneous xenografts of C6 glioma were subjected to tail vein injection of PFP/PFB/DiR-PPEHD, DiR-PPEHD and PFP/PFB/DiR-MBs respectively at a dose of 4 mg DiR per Kg body weight. Mice under isoflurane anesthesia were imaged on an in vivo fluorescence imaging system (Carestream In-Vivo Imaging System FX PRO, USA). At 48 h after nanoprobe injection, all animals were sacrificed and their organs of interest (liver, heart, spleen, lung, kidney) and tumor were excised for ex vivo imaging. To determine the effect of LFUS exposure on drug release in tumor site, release behavior of DiR as a model drug were investigated using the same in vivo fluorescent imaging system. Mice bearing subcutaneous xenografts of C6 glioma received the treatments of PFP/PFB/DiR-PPEHD/LFUS(+),

PFP/PFB/DiR-PPEHD/LFUS(-),

DiR-PPEHD/LFUS(+)

and DiR-PPEHD/LFUS(-), respectively. Sample solutions were injected via tail vein. In treatments with LFUS irradiation, LFUS (1 MHz, 50% duty cycle, acoustic pressure 2.0 MPa) was applied for 5 min in tumor area at 11 h after nanoprobe injection. The fluorescent images were acquired before and 5 min or 2 h after LFUS irradiation. The 720 nm excitation filter was used and the 790 nm emission fluorescence was collected. In vitro and in vivo power Doppler imaging. In vitro ultrasound (US) imaging was performed using clinical US scanning system under power Doppler mode (Acuson Sequoia 512, Siemens, USA). Room temperature was adopted for all studies unless otherwise noted. The PFP/PFB/DOX-PPEHD solutions (1.0 mg mL-1) at pH 7.4, pH 6.8 and pH 6.8 + 45 °C respectively were added into the sample wells in a custom-made 2 % (w/v) agarose gel model to achieve power Doppler imaging using a broadband 15L8-w high-frequency linear transducer at a frequency of 10 MHz and a transmitted power of 18 dB. The focal zone was placed at the center of the sample well with a depth of 3 cm. Horizontal imaging of the probe through the agarose gel model was conducted. DOX-PPEHD without encapsulating PFP/PFB was used as a control. 19


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In the in vivo imaging study, the PFP/PFB/DOX-PPEHD solution (100 µL, 5 mg mL-1) was injected into mice bearing subcutaneous C6 glioma via tail vein at a dose of 2.5 mg DOX per Kg body weight. Power Doppler imaging was obtained at 5 h, 11 h and 24 h after injection using clinical US scanning system (Logiq 9 digital premium ultrasound system, GE, USA) at room temperature and 45 °C, respectively. Tumor area of mice was heated using an infrared heater (TANITA, China), and then the real temperature of tumor area was measured with a thermometer (MGS, China). In vivo US imaging was detected on a broadband ML6-15D high-frequency linear transducer, and 5-10 mm space between the transducer and tumor was filled with ultrasonic transmission gel. The imaging in power Doppler mode was detected using frequency of 7.5 MHz and transmitted power of 18 dB. Drug diffusion inside tumor tissue. Mice bearing subcutaneous C6 glioma divided into different groups were subjected to the treatments of PFP/PFB/DOX-PPEHD/LFUS(+), PFP/PFB/DOX-PPEHD/LFUS(-),

DOX-PPEHD/LFUS(+)

and

DOX-PPEHD/LFUS(-),

respectively. Nanoprobe solutions were injected via tail vein at a dose of 2.5 mg DOX per Kg body weight. LFUS (1 MHz, 50 % duty cycle, acoustic pressure 2.0 MPa) was applied exclusively in tumor area for three times (5 min each) at 11 h after nanoprobe injection. In order to label tumor blood vessels, FITC-lectin was injected via tail vein at 2 h after LFUS exposure of tumor (dose: 10 mg per Kg body weight). The mice were anesthetized and perfused with normal saline to remove the residual fluorescent probes in blood circulation. Finally, tumor tissues from all groups were collected and the frozen sections were observed on CLSM (100×). Tumor growth inhibition. Mice bearing subcutaneous C6 glioma were randomly divided into five groups (n = 6) for treatments with PFP/PFB/DOX-PPEHD/LFUS(+), PFP/PFB/DOX-PPEHD/LFUS(-), DOX/PPEHD/LFUS(+), DOX/PPEHD/LFUS(-) and PBS, respectively. When tumor volume reached about 50 mm3, treatments were performed. 100 µL of sample solution was injected into each nude mouse at 2.5 mg DOX per Kg body weight via 20


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tail vein at the time points of 1, 4, 7 and 10 days, respectively. At 11 h, 17 h and 24 h after injection, tumor site was irradiated with LFUS (1 MHz, 50% duty cycle, acoustic pressure 2.0 MPa, 5 min) if required. Mice in the control group received tail vein injection of 100 µL PBS at the same time points. The treatment lasted for thirty days. Tumor volume was measured every three days with vernier caliper and calculated using the equation: volume = 0.5 × l × w2, where “w” and “l” represent width and length of tumor. The survival rate was caculated by the Kaplan-Meier survival analysis using SPSS statistical software. Histological and immunohistochemical assays. After mice were sacrificed, tumors were collected and fixed for 24 h with freshly prepared 10% PBS buffered formalin. At least five paraffin sections were prepared for each animal. Tissue sections (5 µm) were subjected to Haematoxylin/Eosin (H&E) staining after deparaffinization. For assessment of apoptosis, in situ cell death analysis using detection kit Fluoresein (Roche) was performed according to manufacturer's protocol. The tumor tissue sections were deparaffinized with xylene and ethanol, and then washed with PBS. After sample areas were dried, TUNEL reaction mixture was subsequently added at 37 °C for 60 min. Slides were rinsed with PBS for 3 times, and then nuclei staining with Hochest 33342 was performed. Finally, all samples were analyzed under CLSM. Statistical analysis. Statistical analysises of data were performed using one-factor analysis of variance (SPSS software, version 13.0, SPSS Inc.). The data were expressed as mean ± SD. *P < 0.05 was considered statistically significant. All statistical tests were two-sided.

ASSOCIATED CONTENT The authors declare no competing financial interest. Supporting Information Available: Illustration of the synthetic approach and 1H NMR, FTIR, GPC characterizations of diblock polymer mPEG-PAsp(DEA-co-His-co-DIP), particle 21


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Page 22 of 41

size and TEM images of PFP/PFB/DOX-PPEHD nanoprobes at different conditions, serum stability of PFP/PFB/DOX-PPEHD nanoprobes, in vitro fluorescence release and quantitative analysis of DOX from PFP/PFB/DOX-PPEHD nanoprobes, standard curves of PFP and PFB, in vitro power Doppler imaging at different conditions, in vivo DiR fluorescence imaging and power Doppler imaging, body weight curves of mice in the treatment experiment. This material is available free of charge via the Internet at http://pubs.acs.org.

AUTHOR INFORMATION Corresponding Authors *E-mail: [email protected]. *E-mail: [email protected]. *E-mail: [email protected]. ORCID Xintao Shuai: 0000-0003-4271-0310 Author Contributions The manuscript was written through contributions of all authors. L.Z. and T.Y. contributed equally to this work. L.Z., T.Y., R.Z., K.L. and X.T. conceived the idea, designed the experiments, and wrote the paper. L.Z. and B.L. prepared the probe and performed the majority of the in vitro assays. T.Y., C.Q. and Q.Z. did the in vivo experiments. #

L. Zhang and T. Yin contributed equally to this work.

ACKNOWLEDGMENTS This work was financially supported by the National Natural Science Foundation of China (U1401242,

51225305,

81430038),

National

Basic

Research

Program

of

China

(2015CB755500) and Natural Science Foundation of the Guang dong Province (2014A030312018). 22


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FIGURE LEGENDS Figure 1. Schematic illustration of theranostical nanoprobe showing tunable size and performance in vivo derived from multi-stimulation sensitivity.

Figure 2. Characterization of the pH-sensitive copolymer. (a) The chemical structures at two different pH. (b) 1H NMR spectrum of pH-sensitive copolymer in DMSO-d6. (c) 1H-NMR spectra of mPEG-PAsp(DEA-co-His-co-DIP) in D2O at pD 7.4, pD 6.8 and pD 5.0 (DCl was used to adjust the solution pD).

Figure 3. Test on stepwise stimuli steadily regulate morphology of nanoprobe based on polymersome. (a) Transmission electron microscopic images, (b) schematic illustration of morphology and (c) in vitro power Doppler imaging of PFP/PFB/DOX-PPEHD nanoprobe at different conditions (nanoprobe concentrations: 1 mg mL-1, all samples were stained with uranyl acetate for TEM analysis. (c) power Doppler imaging: 15L8-w broadband high-frequency linear transducer, frequency of 10 MHz and a transmitted power of 18 dB, white arrows and circles indicated signal region). (d) DOX fluorescence intensities and (e) particle size change of PFP/PFB/DOX-PPEHD nanoprobe aqueous solutions at different conditions, and all measurements were conducted 2 h after condition adjustment (nanoprobe concentration: 0.5 mg mL-1. **P < 0.01, *P < 0.05. LFUS if applied: 1 MHz, 50% duty cycle, acoustic pressure 2.0 MPa, 5 min. Data were mean ± SD, n = 3).

Figure 4. Nanoprobe-enhanced tumor imaging in vivo. (a) DiR fluorescence imaging showing tumor accumulation of PFP/PFB/DiR-PPEHD nanoprobe at different time points after tail vein injection into nude mice bearing C6 glioma tumor. Tumors are circled with white dashed lines. (b) Tumor-focusing power Doppler US imaging of animals after tail vein injection of 32


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PFP/PFB/DOX-PPEHD nanoprobe at 37 °C and 45 °C. Power Doppler mode: broadband ML6-15D linear transducer, frequency of 7.5 MHz and transmitted power of 18 dB. The squares and arrows indicated the imaging zone highlighting tumor. DiR dose in (a): 4 mg per Kg body weight, DOX dose in (b): 2.5 mg per Kg body weight.

Figure 5. LFUS-triggered intratumor drug release and tumor-penetrating drug delivery. (a) Schematic illustration of LFUS promoting intratumor drug release and penetration. (b) In vivo DiR fluorescence imaging and (c) quantitative analysis of DiR fluorescence intensities of tumors with or without LFUS exposure at different time points pre- and after exposure. LFUS exposure if required was applied at 11 h after tail vein administration of DiR-containing nanoprobes. (d) Confocal laser scanning microscopic observation of frozen sections of C6 glioma from animals receiving DOX-containing nanoprobes. Dose: 2.5 mg DOX per Kg body weight. scale bar: 50 µm. LFUS irradiation (1 MHz, 50% duty cycle, acoustic pressure 2.0 MPa, 3 times, 5 min each) of tumor was applied at 11 h after injection, and tumor tissue was obtained at 2 h after LFUS irradiation.

Figure 6. Antitumor activity of nanoprobe improved by LFUS. (a) Schematic illustration of PFP/PFB/DOX-PPEHD and low frequency ultrasound (LFUS) combination therapy to inhibit tumor growth. (b) Tumor growth inhibition and (c) cumulative survival of nude mice bearing C6 glioma after tail vein administration of different therapeutic formulations (n=6, dose per injection: 2.5 mg DOX per Kg body weight, four times injections indicated by dashed arrows were performed with an interval of 3 days). *P < 0.05 vs. PFP/PFB/DOX-PPEHD/LFUS(-), DOX-PPEHD/LFUS(+) and DOX-PPEHD/LFUS(-). The values in (b) are mean ± SD. (d) Histological H&E analyses of C6 glioma sections at 30 d after the first treatment (blue color: 33


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nuclei, red color: extracellular matrix and cytoplasm). Cell apoptosis examined by fluorescent TUNEL assay via CLSM. Blue fluorescence indicated nuclei stained with Hoechst 33342, and green fluorescence indicated apoptotic nuclei stained with fluorescein isothiocyanate. Scale bars: 50 µm. LFUS if applied: 1 MHz, 50% duty cycle, acoustic pressure 2.0 MPa, 3 times and 5 min each.

Table 1. Characteristics of the block copolymers. Sample

Mn a (Da)

Mn b (Da)

Mw/Mn b

18400

20600

1.05

22800

24000

1.08

mPEG-PBLA mPEG-PAsp(DEA-co-His-co-DIP) a)

1

b)

Calculated from H NMR spectra, determined by GPC analyses using PEG as a calibration standard.

Table 2. Particle size of PFP/PFB/DOX-PPEHD at different temperatures. All measurements were conducted 30 min after condition adjustment. Data were shown as mean ± SD (n = 3). Temperature (°C)

25

37

40

45

50

55

Size at pH 7.4 (nm)

172±6

178±11

267±9

415±10

469±17

384±28

Size at pH 6.8 (nm)

426±16

437±22

485±19

509±18

432±34

362±47

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Figure 1 162x147mm (300 x 300 DPI)


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Figure 2 159x146mm (600 x 600 DPI)


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Figure 3 231x319mm (300 x 300 DPI)


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Figure 4 109x71mm (600 x 600 DPI)


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Figure 5 191x201mm (300 x 300 DPI)


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Figure 6 188x198mm (300 x 300 DPI)


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Table of Contents Graphic 62x39mm (300 x 300 DPI)


Size-Modulable Nanoprobe for High-Performance Ultrasound Imaging

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