- Nano Express
- Open Access
Single-step assembly of polymer-lipid hybrid nanoparticles for mitomycin C delivery
- Yunfeng Yi†1,
- Yang Li†2, 3,
- Hongjie Wu4,
- Mengmeng Jia2,
- Xiangrui Yang2,
- Heng Wei2,
- Jinyan Lin2,
- Shichao Wu2, 3,
- Yu Huang2,
- Zhenqing Hou2 and
- Liya Xie5Email author
© Yi et al.; licensee Springer. 2014
- Received: 16 August 2014
- Accepted: 25 September 2014
- Published: 8 October 2014
Mitomycin C is one of the most effective chemotherapeutic agents for a wide spectrum of cancers, but its clinical use is still hindered by the mitomycin C (MMC) delivery systems. In this study, the MMC-loaded polymer-lipid hybrid nanoparticles (NPs) were prepared by a single-step assembly (ACS Nano 2012, 6:4955 to 4965) of MMC-soybean phosphatidyhlcholine (SPC) complex (Mol. Pharmaceutics 2013, 10:90 to 101) and biodegradable polylactic acid (PLA) polymers for intravenous MMC delivery. The advantage of the MMC-SPC complex on the polymer-lipid hybrid NPs was that MMC-SPC was used as a structural element to offer the integrity of the hybrid NPs, served as a drug preparation to increase the effectiveness and safety and control the release of MMC, and acted as an emulsifier to facilitate and stabilize the formation. Compared to the PLA NPs/MMC, the PLA NPs/MMC-SPC showed a significant accumulation of MMC in the nuclei as the action site of MMC. The PLA NPs/MMC-SPC also exhibited a significantly higher anticancer effect compared to the PLA NPs/MMC or free MMC injection in vitro and in vivo. These results suggested that the MMC-loaded polymer-lipid hybrid NPs might be useful and efficient drug delivery systems for widening the therapeutic window of MMC and bringing the clinical use of MMC one step closer to reality.
Nanotechnology represents a powerful tool in the field of medicine to deal with cancer[1, 2], which is still a leading cause of death worldwide. Nevertheless, there is the limited availability of nanotechnology-based drug formulation which is approved by FDA in clinical use. It is a formidable challenge to develop the smarter and better nanoscaled drug delivery systems. Up to now, both lipid-based and polymer-based nanoscaled drug delivery systems have represented two dominant classes of nanoscaled drug delivery systems for anticancer drug delivery and shown promising efficacy for clinical treatment[4–6]. Liposome, one of the most investigated lipid-based carriers, is a spherical vesicle composed of lipid material and has been widely used for drug delivery due to improved safety profile, high delivery efficiency, favorable pharmacokinetics, and ease of surface modification. However, their clinical uses by intravenous injection are significantly limited because of the insufficient drug loading, burst drug release, instability, and toxicity. Polymeric nanoparticles (NPs), one of the most investigated polymer-based carriers, have small particle size, ability to load drug of poor solubility and permeability, appropriate physiological stability, well-storage stability, controlled and sustained drug release, and long systemic circulation time[5, 8]. In addition, the preparation of polymeric NPs by nanoprecipitation or emulsion is simple and scalable. However, the biocompatibility of the polymeric NPs formed by most synthetic polymers is not as high as the liposomes, especially at the cellular and animal level. Thus, it is necessary to develop a novel and practical drug carrier which can combine the desired advantages and avoid the undesired disadvantages of both the lipid-based and polymer-based carriers for clinical application.
In recent years, the polymer-lipid hybrid NPs designed to merge the best of both worlds had been reported in several groups. The first was that the polymer-lipid hybrid NPs were prepared via the fusion of the polymeric NPs and liposomes. The preparation of the polymer-lipid core-shell structure usually required a two-step process: the initial formation of polymeric NPs and subsequent encapsulation of the liposomes[8, 12]. However, the process resulted in the technical complexity and the poor control over the final characteristics of the nanoscaled drug delivery systems[8, 13]. The second was that the polymer-lipid hybrid NPs were prepared by a single step which combined the nanoprecipitation (or emulsion) and self-assembly method[14, 15]. The strategy satisfied the need for the development of precise and predictable hybrid nanoscaled drug delivery systems, which would be convenient for future scaling-up.
Mitomycin C (MMC) is a water-soluble anticancer and antibiotic agent by inhibiting DNA synthesis and nuclear division and has been extensively used to treat various cancer including stomach, liver, breast, pancreas, colon, and bladder cancers. MMC was a very poor substrate for P-glycoprotein and retained activity against many types of P-glycoprotein-mediated multidrug resistant tumor cells[17, 18]. However, a major concern of MMC was in the narrow therapeutic index and serious side effects such as severe myelosuppression and gastrointestinal complications[19, 20]. Recently, the drug-phospholipid complex had received significant attention[21, 22] because of the significant improvement of drug efficacy and safety[23–30]. To address the issue, we prepared the MMC-soybean phosphatidyhlcholine complex (MMC-SPC complex) to increase the effectiveness of MMC. On the one hand, MMC interacted with soybean phosphatidylcholine (SPC) via the non-covalent bonds and resulted in almost 100% complexation efficiency. On the other hand, the amphiphilic character of the drug-phospholipid complex could promote the passive transport from the water environment to the lipid-rich cell membrane to increase the drug permeability and enhance the drug bioavailability and improve the drug effectiveness.
All chemical reagents were of analytical grade and used without further purification unless otherwise stated. Mitomycin C (MMC; purity grade = 99.5%) was purchased from Hisun Pharmaceutical Co., Ltd. (Zhengjiang, China). SPC was provided by Lipoid GmbH (Ludwigshafen, Germany). Poly(D,L-lactide) (PLA; 10 kDa) was provided by Daigang BIO Engineer Co., Ltd. (Shandong, China). N,N-dimethylformamide (DMF), 3-(2-benzothiazolyl)-N,N-diethylumbelliferylamine (coumarin-6), fluorescein 5-isothiocyanate (FITC), and Dulbecco's modified Eagle's medium (DMEM) were purchased from Sigma-Aldrich (St. Louis, MO, USA). Iodide [1,1′-dioctadecyl-3,3,3′,3′-tetramethylindotricarbocyanine iodide (DiR) was provided by Molecular Probes Inc. (Eugene, OR, USA). A dialysis bag (Mw = 8, 000 to 14, 000 Da) was ordered from Greenbird Inc. (Shanghai, China). Deionized (DI) water was used throughout. Fetal bovine serum (FBS) was purchased from Gibco Life Technologies (AG, Zug, Switzerland). Trypsin-EDTA (0.25%) and penicillin-streptomycin solution were from Invitrogen (Life Technologies, Basel, Switzerland). All solvents used in this study were high-performance liquid chromatography (HPLC) grade. Mouse hepatoma cell line H22 cells were provided by the Cell Bank at the China Academy of Science.
Preparation of PLA NPs/MMC-SPC
Firstly, MMC and SPC were added to tetrahydrofuran by vigorous agitation in a 40°C water bath for 4 h. The organic solvent was evaporated under reduced pressure. The MMC-SPC loaded PLA NPs (PLA NPs/MMC-SPC) were prepared by a facile dialysis method. Twenty milligrams of PLA and the weighed amount of MMC-SPC (the weight ratio of MMC and SPC was 1:2) were dissolved in 10 mL of DMF accompanied by vigorous stirring, and then, the resulting organic phase was introduced into a dialysis bag. Subsequently, the dialysis bag was placed into 1,000 mL of DI water as the aqueous phase under moderate agitation. The organic phase was dialyzed against the aqueous phase for 8 h, after that, the suspensions were filtered through a 450 and 200 nm polycarbonate membrane (Millipore, Bedford, MA, USA), respectively. Then, the PLA NPs/MMC-SPC were collected by centrifugation with a JA-20 rotor (Beckman Coulter, Inc., Fullerton, CA, USA) at 15,000 rpm at 4°C for 10 min, after that, the PLA NPs/MMC-SPC were resuspended in DI water. The PLA NPs/MMC were prepared by a double emulsion solvent-evaporation technique for comparison. PLA was added into 10 mL of DMF to form the oil phase, and MMC was added into 5 mL of DI water to form the inner water phase. The inner water phase was mixed with the oil phase under stirring and then, the mixture was sonicated in an ice water bath to form a primary emulsion. The primary emulsion was added into DI water with 2% (w/v) PVA and homogenized at of 1,000 rpm.
In addition, the micron-sized PLA NPs/MMC-SPC (used for confocal laser scanning microscopy image) could be formulated as follow: FITC conjugated MMC (FITC-MMC; FITC-MMC was synthesized based on the reaction between MMC and FTIC via a thiourea linkage) was prepared for use as a fluorescence probe. And then, FITC-MMC-SPC complex was prepared in a same way as described above. Thirty-six milligrams of FITC-MMC-SPC complex and 200 mg of PLA were dissolved in 10 mL of DMF, accompanied by vigorous stirring. The mixture was dialyzed against 1,000 mL of DI water for 8 h. Then, the micron-sized PLA NPs/FITC-MMC-SPC were collected by centrifugation at 15,000 rpm at 4°C for 10 min, after that, the micron-sized PLA NPs/FITC-MMC-SPC were resuspended in DI water.
Drug encapsulation efficiency and drug loading content
Characterization of PLA NPs/MMC-SPC
The average particle size, polydispersity index (PDI), and zeta potential of the suspension of the PLA NPs/MMC-SPC were performed by dynamic light scattering (DLS) using a Malvern Zetasizer Nano-ZS (Malvern Instruments, Worcestershire, UK). Particle size was evaluated by intensity distribution. The morphology of the PLA NPs/MMC-SPC was visualized by SEM (LEO 1530VP, Oberkochen, Germany) operating at 20 kV and TEM (JEM 1400, JEOL, Tokyo, Japan) operating at 200 kV.
In vitro stability tests
The lyophilized PLA NPs/MMC-SPC were suspended in phosphate buffer solution (PBS) (pH 7.4) and incubated at 37°C for 48 h. The particle size was assayed at 12 h intervals by DLS.
In vitro drug release
The release of MMC from the PLA NPs/MMC-SPC was determined by a dialysis technique. The lyophilized NPs were dispersed in 3 mL of PBS (1/15 M, pH 7.4) buffer solution. The dispersions were transferred to a dialysis bag (Mw = 8,000 to 14,000 Da) and then subjected to dialysis against 10 mL of PBS at 37°C. At a predesigned time interval, all of the PBS buffer solution was withdrawn and subsequently replaced with the 10 mL of fresh PBS after each sampling. The PLA NPs/MMC and free MMC were used for comparison. The release of MMC was determined by a HPLC method as described above.
H22 cells (mouse hepatoma cell line) were cultured in DMEM supplemented with 10% FBS and 1% penicillin-streptomycin. The cells were cultivated in a humidified atmosphere containing 5% CO2 at 37°C.
In vitro cellular uptake
To facilitate the observation of cellular uptake, coumarin-6 was used as a hydrophobic fluorescence probe to load within the PLA NPs/MMC-SPC and PLA NPs/MMC, respectively (designed as the coumarin-6-PLA NPs/MMC-SPC and coumarin-6-PLA NPs/MMC, respectively). H22 cells were seeded at a density of 1 × 105 cells per well in 6-well plates with their specific cell culture medium. The cells were incubated at 37°C and 5% CO2 for 24 h. One-hundred microliters of the coumarin-6-PLA NPs/MMC-SPC and coumarin-6-PLA NPs/MMC at the equivalent coumarin-6 concentration was added and incubated further for 6 h. The cells were washed with PBS, fixed with 4% paraformaldehyde, and stained with Hoechst 33258 (Sigma-Aldrich, St. Louis, MO, USA). The cells were observed using a Leica TCS SP5 confocal laser scanning microscopy (Leica Microsystems, Mannheim, Germany).
To further quantitatively investigate the cellular uptake, coumarin-6 was used as a fluorescence probe to load within the PLA NPs/MMC-SPC and PLA NPs/MMC, respectively (designed as the coumarin-6-PLA NPs/MMC-SPC and coumarin-6-PLA NPs/MMC, respectively). H22 cells were seeded in 6-well plates with a density of 2 × 105 cells/mL and incubated for 24 h, and then, the original medium was replaced with the coumarin-6-PLA NPs/MMC-SPC and coumarin-6-PLA NPs/MMC-SPC at the equivalent coumarin-6 concentration. The cells were incubated for the predesigned time at 37°C and then washed with cold PBS and harvested by 0.25% trypsin-EDTA. The harvested cells were suspended in PBS and centrifuged at 1,000 rpm for 5 min at 4°C. After washing and centrifugation, the cells were resuspended in PBS and performed by a Beckman Coulter EPICS XL flow cytometer (Beckman Coulter, Fullerton, CA, USA).
Intracellular drug delivery
To understand the intracellular distribution of the drug inside H22 cells, FITC conjugated MMC (FITC-MMC; FITC-MMC was synthesized based on the reaction between MMC and FTIC via a thiourea linkage) was prepared for use as a fluorescence probe. H22 cells were incubated for 24 h and then cultured with the PLA NPs/FITC-MMC-SPC or PLA NPs/FITC-MMC for 12 h at 37°C. The cells were imaged using a confocal laser scanning microscopy. Rhodamine phalloidin was used to stain cytoskeleton. Hoechst 33258 (Sigma-Aldrich, St. Louis, MO, USA) was used to stain nuclei.
In vitro cell viability
The cytotoxicity of the drug delivery systems was measured using a MTT assay (Sigma-Aldrich, St. Louis, MO, USA) according to the manufacturer's suggested procedures. H22 cells were exposed to the PLA NPs/MMC-SPC, PLA NPs/MMC or free MMC with different MMC concentrations for 24 h. The data were expressed as the percentage of surviving cells.
In vivo blood circulation
All the animal procedures complied with the guidelines of the Xiamen University Institutional Animal Care and Use Committee. The experiments were performed on adult male SD rat weighing 200 ± 20 g (mean ± SD) from Shanghai Laboratory Animal Center. DiR was used as a near-infrared fluorescence probe to load within the PLA NPs/MMC-SPC and PLA NPs/MMC, respectively. One-hundred microliters of the DiR-PLA NPs/MMC-SPC and DiR-PLA NPs/MMC at the equivalent DiR concentration was respectively injected intravenously through the tail vein of rats. At timed intervals, 200 μL of blood was collected and stored in heparin containing eppendorf tubes at 4°C for further analysis. The plasma was separated from the blood by centrifuging at 2,000 rpm for 5 min. The plasma was diluted with methanol and centrifuged to remove the insoluble solid. At the excitation wavelength of 748 nm, the fluorescence intensity at 780 nm was measured using a microplate reader and the corresponding DiR concentration was calculated according to an established standard curve.
In vivo fluorescence imaging
For in vivo imaging, 0.2 mL of the DiR-PLA NPs/MMC-SPC was injected into the nude mice bearing H22 tumor via the lateral tail vein. Imaging was performed at 1, 3, and 12 h after injection using a Maestro™ in vivo imaging system (Cambridge Research & Instrumentation, Woburn, MA, USA). At 12 h post-injection, the mice were sacrificed. The tumor and normal tissues (tumor, liver, spleen, lung, kidney, and heart) were excised, followed by washing the surface with 0.9% NaCl for the ex vivo imaging of DiR fluorescence using a Maestro™ in vivo imaging system (Cambridge Research & Instrumentation, Woburn, MA, USA). The mice treated without 0.2 mL of the DiR-PLA NPs/MMC-SPC were used for comparison.
In vivo anticancer effect
Kunming mice aged 4 to 5 weeks (clean class, 18 to 22 g) were supplied by Xiamen University Laboratory Animal Center and used in this study. Subcutaneous tumors were established in the mice by subcutaneous inoculation of 5 × 106 H22 cells in the right axillary region of the mice before the treatment. H22 tumors were not hormone dependent and easily grown in the subcutaneous layer of mice. The tumors were allowed to grow for 1 week after tumor transplantation, after that, the treatment was performed. The H22 tumor-bearing mice were randomly divided into 4 groups (10 mice per group): group 1 for 0.9% NaCl, group 2 for PLA NPs/SPC, group 3 for free MMC injection, and group 4 for PLA NPs/MMC-SPC. The mice were intravenously administrated at 4 mg/kg (equivalent MMC concentration) every 2 days for 3 times. Each mouse was earmarked and followed individually throughout the whole experiments. The tumor size and body weight were then monitored every 2 days. The mice were euthanized on day 12. The greatest longitudinal diameter (length) and the greatest transverse diameter (width) of the tumor were measured using a vernier caliper, and the tumor volume was calculated using length × width2 × 0.5.
Results and discussion
Preparation and characterization of PLA NPs/MMC-SPC
The hybrid PLA NPs/MMC-SPC comprised a hydrophobic PLA core and an amphiphilic MMC-SPC shell. Once intravenously administrated, the polymer-lipid hybrid NPs were accumulated at the tumor site via the enhanced permeability and retention (EPR) effect. And then, the polymer-lipid hybrid NPs were internalized by the cells via endocytosis. Lastly, MMC was released from the polymer-lipid hybrid NPs and delivered to the nuclei, presenting the anticancer activity.
The effect of the MMC-SPC concentration on the size, zeta potential, and encapsulation efficiency of the hybrid PLA NPs/MMC-SPC was investigated. As shown in Additional file1: Figure S1 in supporting information, when the concentration of MMC-SPC increased, the particle size of the hybrid PLA NPs/MMC-SPC initially decreased and then increased. It was believed that at the appropriate MMC-SPC concentration (0.36 mg/mL), the amount of the amphiphilic MMC-SPC complex was in the range to able to cover the surface of the hydrophobic PLA core. However, when the MMC-SPC concentration was too high, the excess of the MMC-SPC complex might participate in the self-organization of the large size, vesicle-like structures (50 to 500 nm)[7, 22]. Therefore, in the following studies, a MMC-SPC concentration of 0.36 mg/mL was used to prepare the hybrid PLA NPs/MMC-SPC.
Particle size, PDI, zeta potential, drug encapsulation efficiency, and drug loading content of the PLA NPs/MMC-SPC and PLA NPs/MMC
Particle size (nm)
Zeta potential (mV)
Drug encapsulation efficiency (%)
Drug loading content (%)
205.3 ± 5.7
0.110 ± 0.021
-36.3 ± 2.9
37.6 ± 2.9
10.3 ± 2.4
327.1 ± 12.9
0.320 ± 0.047
-18.7 ± 1.5
34.5 ± 3.8
1.7 ± 0.4
As it is reported, a particle size (200 to 300 nm) may be suited for the EPR effect[39, 40], providing a passive targeting. A high zeta potential (lower than -30 mV or higher than 30 mV) can provide an electrostatic repulsion to avoid the aggregation of the particles. A high drug loading content is essential for intravenous administration, which will reduce the dosage of carriers and the possible side effects to the patients[42, 43]. Thus, SPC had much higher emulsification efficiency than PVA. Moreover, SPC was a natural product, and the PVA was a chemical. The former can thus cause less side effects than the latter. All of the results demonstrated a prospect of the useful nanoscaled drug delivery systems (PLA NPs/MMC-SPC) with a nanoscaled particle size, a narrow particle size distribution (Figure 2A and Table 1), a high zeta potential (Figure 2B and Table 1), a spherical shape (Figure 2C,D), a well-stability (Figure 2E, discussed as below) and a high drug loading content (Table 1) for drug delivery.
In vitro stability
The clinical use of the PLA NPs as intravenous drug delivery systems is limited due to their poor physiological stability. We thus investigated if the hybrid PLA NPs/MMC-SPC can overcome this drawback. As shown in Figure 2E, the hybrid PLA NPs/MMC-SPC exhibited a structural stability under physiological conditions as demonstrated by no significant particle size change over 48 h. We speculated that the presence of the phospholipid complex on the surface of the hybrid PLA NPs/MMC-SPC prevented the aggregation of the nanoscaled drug delivery systems.
In vitro drug release
In vitro cellular uptake
To more precisely understand how the composition of formulations affected the cell internalization of the NPs, H22 cells were incubated with the coumarin-6-PLA NPs/MMC-SPC and coumarin-6-PLA NPs/MMC and subsequently analyzed the cells by flow cytometry. As clearly shown in Figure 4B, both the coumarin-6-PLA NPs/MMC-SPC and coumarin-6-PLA NPs/MMC indeed showed a time-dependent cell uptake. The cellular uptake of the coumarin-6-PLA NPs/MMC-SPC was not significantly different from that of the coumarin-6-PLA NPs/MMC-SPC after incubation of 0.5 h. In contrast, the cellular uptake of the coumarin-6-PLA NPs/MMC-SPC was significantly higher than that of the coumarin-6-PLA NPs/MMC after incubation of 1.5 and 4.5 h. The different phenomenon could be explained by cell penetration rate of the nanoscaled drug delivery systems depending on the NP concentration differences between the internal and external environment of the cell membrane or the interaction of the NPs and the cell membrane. It should be concluded that in the case of the incubation of a short period of time, the NP concentration differences between the internal and external environment of the cell membrane played a main role in the cell penetration rate of the nanoscaled drug delivery systems; however, in the case of the incubation of a long period of time, the interaction of the NPs and the cell membrane played a vital role.
Consistent with the observations by confocal laser scanning microscopy, the result of flow cytometry confirming that enhanced cellular internalization (45.6% increased for 4.5 h) of the PLA NPs/MMC-SPC compared with the PLA NPs/MMC. The lipophilicity and liposolubility of SPC on the surface of the hybrid PLA NPs/MMC-SPC perhaps increased the endocytosis and facilitated the passive delivery of the PLA NPs/MMC-SPC NPs to the interior of the cells.
Intracellular drug delivery
For both the PLA NPs/FITC-MMC-SPC and PLA NPs/FITC-MMC, the green fluorescence was presented in the cell. This finding was similar to the reported work. It was also reported that the premature released FITC or FITC-MMC could be not internalized by the cells. To the result, we proposed two possible reasons, one was that the intracellular drug release and transport played a dominate role. A part of drug was released from the NPs and transported inside the cells. Alternatively, the other one was that the intracellular drug accumulation and diffusion played a main role. A high nuclear accumulation of FITC-labeled MMC resulted in the subsequent diffusion to other cytoplasmic areas. Regardless of either mechanism, the result confirmed that the drug delivered to the nuclei resulted from the effective cellular uptake, followed by the efficient intracellular internalization and accumulation of drug.
It should be noted that, the fluorescence signals of H22 cells incubated with the PLA NPs/FITC-MMC-SPC were significantly increased compared with those incubated with the PLA NPs/FITC-MMC at the same incubation time under the identical instrumental conditions, the result combined with the in vitro cellular uptake further confirmed that the interaction of SPC on the surface of the hybrid PLA NPs/FITC-MMC-SPC, and the cell membrane enhanced the cellular uptake of the nanoscaled drug delivery systems, hence, increased the intracellular internalization and accumulation of FITC-labeled MMC. A sufficient intracellular drug concentration was essential for an enhanced anticancer activity. Escape of the delivered MMC from endo/lysosome into nuclei was important as the target site of MMC was nuclear DNA and MMC could interrupt its function. Although more studies were needed to better understand the nuclear radiolabeled MMC delivery and the underlying mechanism, the result did suggest that our NPs were capable of effectively mediating intracellular drug delivery.
In vitro cell viability
Besides, the cytotoxicity PLA NPs/MMC-SPC were not as effective as the free MMC against H22 cells at an equivalent MMC concentration (Figure 6), which was most probably because of the prolonged drug release from the hybrid drug-loaded NPs (see Figure 3). Such a sustained and prolonged drug release of the PLA NPs/MMC-SPC might result in the progressive increase of intracellular drug concentration for cell death. On the contrary, the free drugs can be rapidly transported into cells by passive diffuse owing to the driving force of a pH and concentration gradient and directly inhibit the cell growth without the drug release. Although the PLA NPs/MMC-SPC induced a decline in the cytotoxicity of MMC, the introduction of the nanoscaled drug delivery systems could increase the passive targeting efficiency in vivo (discussed below). Once accumulated at the tumor site, the polymer-lipid hybrid NPs as a whole would easily into the interior of the tumor cells to exert the pharmacological effects.
In vivo stability
Ex vivo fluorescence imaging
In vivo anticancer effect
Next, we examined the anticancer effect of the PLA NPs/MMC-SPC after the accumulation at the tumor site. As shown in Figure 8C, the intravenous injection of the PLA NPs/MMC-SPC and free MMC inhibited the tumor growth, whereas the MMC-free PLA NPs/SPC did not slow the tumor growth. The result showed that MMC, the delivery of which was mediated by the hybrid PLA NPs/MMC-SPC, was responsible for the tumor growth inhibition. More importantly, the tumor growth of mice treated with the PLA NPs/MMC-SPC was much slower than that of mice treated with the free MMC, and the difference became more significant after day 12. In addition, the loss of body weight in mice accompanied the treatment with the free MMC in this study (Figure 8D) but was not found in the treatment with the hybrid PLA NPs/MMC-SPC. All of the results suggested that compared with the free MMC, the hybrid PLA NPs/MMC-SPC showed a significantly enhanced therapeutic efficacy while reducing the side effect of chemotherapy drug.
We concluded that the following reasons might be involved. Firstly, the free MMC is cleared too rapidly, and thus, the low concentrations of drugs in the tumor tissues will result in suboptimal therapeutic effects. On the contrary, the EPR effect and the sustained release of the nanoscaled drug delivery systems may result in a sufficient intracellular drug. Secondly, the nanoscaled drug delivery systems can help the loaded drug effectively enters the interior of cells by endocytosis. Lastly, the lipophilicity and liposolubility of SPC on the surface of the polymer-lipid hybrid NPs efficiently help the NPs transport from the surrounding water-soluble environment to the lipid-rich cell membrane and enter the internal environment of the cells, leading to the increased internalization and accumulation of drug inside the cells. Therefore, the present study suggested that after the intravenous administration, the hybrid PLA NPs/MMC-SPC were useful in significantly improving the anticancer effect of MMC while reducing its toxicity compared to the MMC injection for clinical treatment.
We have developed the MMC-loaded polymer-lipid hybrid NPs for sustained and controlled release of MMC by a single-step self-assembly. The composition of the PLA NPs/MMC-SPC not only affected their drug release but also influenced their cellular uptake and anticancer efficacy. We concluded that natural SPC had great advantages over traditional PVA with less burst drug release, more cellular uptake, and higher anticancer efficacy. The PLA NPs/MMC-SPC might efficiently ensure the nuclear delivery of MMC to induce cell death. More importantly, the PLA NPs/MMC-SPC showed the improved therapeutic efficiency compared with the free MMC injection. All of the results suggested that the MMC-loaded polymer-lipid hybrid NPs have promising potential as attractive and practical nanoscaled MMC delivery systems for cancer therapy.
Yunfeng Yi and Yang Li are co-first authors.
This work was founded by The Medical Science and Technology Innovation Project of Nanjing Military Command (10MA078, 2010) and Xiamen Science and Technology Project (2010S0862).
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