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Polymeric nanoparticles (NPs) and dendrimers are two major classes of nanomaterials that have demonstrated great potential for targeted drug delivery. However, their targeting efficacy has not yet met clinical needs largely because of a lack of control over their targeting kinetics, which often results in rapid clearance and off-target drug delivery. To address this issue, we have designed a novel hybrid NP (nanohybrid) platform that allows targeting kinetics to be effectively controlled through hybridization of targeted dendrimers with polymeric NPs. Folate (FA)-targeted generation 4 poly(amidoamine) dendrimers were encapsulated into poly(ethylene glycol)-b-poly(D,L-lactide) (PEG-PLA) NPs using a double emulsion method, forming nanohybrids with a uniform size (~100 nm in diameter) at high encapsulation efficiencies (69–85%). Targeted dendrimers encapsulated within the NPs selectively interacted with FA receptor (FR)-overexpressing KB cells upon release in a temporally controlled manner. The targeting kinetics of the nanohybrids were modulated using three different molecular weights (MW) of the PLA block (23, 30, and 45 kDa). The release rates of the dendrimers from the nanohybrids were inversely proportional to the MW of the PLA block, which dictated their binding and internalization kinetics with KB cells. Our results provide evidence that selective cellular interactions can be kinetically controlled by the nanohybrid design, which can potentially enhance targeting efficacy of nanocarriers.
Polymeric nanocarriers have been widely investigated as a versatile platform for controlled drug delivery to target tissues.1–3 Among many polymeric materials, poly(amidoamine) (PAMAM) dendrimers hold great promise due to their well-controlled structure and multifunctionality.4–6 Through conjugation of targeting ligands, dendritic nanodevices (5–10 nm in diameter) have been shown to be effective in achieving selective tumor targeting through specific ligand-receptor interactions, i.e., active targeting.7–9 On the other hand, larger nanomaterials such as polymeric nanoparticles (NPs) and liposomes are well suited to exploit another targeting strategy termed passive targeting. By utilizing the characteristic tumor biology highlighted by the enhanced permeability and retention (EPR) effect,10–11 these size-controlled nanocarriers (typically 50–200 nm in diameter) have shown selective accumulation in tumor sites.12–13 Each of the two targeting strategies, however, suffers from several limitations.14–15 Despite the enhanced active targeting efficacy of dendrimers due to their increased molecular flexibility that facilitates the multivalent binding effect,16–17 their small size limits their passive targeting capability.18 Targeted dendrimers with surface-conjugated folic acid (FA) have also been associated with off-target delivery and rapid clearance in vivo.19 In contrast, the relatively large size of the NPs hinders their effective penetration into tumor sites due to limited diffusivity.20 The rigidity of the actively targeted NPs compared to the flexible dendrimers can also prevent the full utilization of the multivalent binding effect, leading to much lower binding avidities (~100-fold enhancement over free FA)20 compared to the dendritic nanodevices (up to 170,000-fold enhancement).16, 21
Nanocarriers that combine passive targeting based on size control and active targeting through surface-immobilized ligands have been reported for a variety of systems including liposomes,22 micelles,23 and polymeric NPs.24 However, these approaches do not fully address the issues of decreased circulation times and rapid clearance. For example, studies have shown that the prolonged circulation times achieved by PEGylated nanocarriers are compromised by the addition of targeting ligands to the outer surface.25–27 These systems also lack the control over the targeting kinetics, due to the surface-exposed targeting ligands. To address these issues, we have designed novel multi-scale hybrid NP systems, or nanohybrids, that combine FA-targeted dendrimer conjugates with polymeric NPs to exploit the strengths and to address the limitations of each individual nanocarrier. This system is based on our previously reported hybrid NPs where poly(ethylenimine)-rhodamine (PEI-RHO) conjugates were encapsulated within protective outer layers of biodegradable polymer-based NPs or biocompatible liposomes.28 This hybrid design is unique in that the biologically active polymer conjugates (PEI-RHO) are protected by an outer shell of biodegradable polymers, or biocompatible lipids, allowing precise control over the cellular interaction kinetics of the bioactive polymers.
In this study, the FA-targeted generation 4 (G4) PAMAM dendrimers were encapsulated into biodegradable poly(ethylene glycol)-b-poly(D,L-lactide) (PEG-PLA) NPs to produce nanohybrids. Through this design, we present a proof-of-concept study where kinetic control over the selective interactions of FA-targeted dendrimers with folate receptor (FR)-overexpressing KB cells (KB FR+) was achieved in vitro. The design rationale for these nanohybrids is to ultimately achieve sequential utilization of passive and active targeting, i.e., passive accumulation at the tumor site by the controlled size of the nanohybrids, followed by active targeting to individual tumor cells by the dendrimers upon their release from the NPs. This paper progresses by testing three hypotheses: 1) multifunctional G4 PAMAM dendrimers can be successfully encapsulated into biodegradable PEG-PLA copolymers to form nanohybrids; 2) the cellular interaction and targeting kinetics of FA-targeted dendrimers can be temporally controlled through the nanohybrid platform; and 3) the targeting kinetics can be further modulated by controlling the molecular weight (MW) of the biodegradable encapsulating copolymers. Here we report, for the first time to our knowledge, a nanohybrid design that presents a promising delivery platform to enable precise control over its cellular targeting and release kinetics, which has the great potential to overcome the limitations of the existing nanocarrier systems.
Generation 4 (G4) PAMAM dendrimer, N-hydroxysuccinimide-rhodamine B (NHS-RHO), folic acid (FA), N-(3-dimethylaminopropyl)-N′-ethylcarbodiimide hydrochloride (EDC), N-hydroxysuccinimide (NHS), glycidol, tin(II)2-ethylhexanoate, poly(ethylene glycol) monomethyl ether (mPEG) (MW 5,000 Da), poly(vinyl alcohol) (PVA, 87–89% hydrolyzed, MW 13,000–23,000 Da), and dichloromethane (DCM), were all obtained from Sigma-Aldrich (St. Louis, MO). D,L-lactide was purchased from Polysciences Inc. (Warrington, PA). Poly(ethylene glycol)-b-poly(lactide) (MW 5,000-b-23,000) was obtained from Polymer Source (Quebec, Canada). All other chemicals used in this study were purchased from Sigma-Aldrich unless specified otherwise.
G4 PAMAM dendrimers were fluorescently labeled by conjugation with NHS-RHO as described earlier.28 Amine-terminated G4 (G4-NH2, 30 mg, 2.0×10−6 mol) was dissolved in 4 mL sodium bicarbonate buffer (pH 9.0), to which 500 μL of NHS-RHO (1.1×10−5 mol) in DMSO was added, and the reaction mixture was vigorously stirred at room temperature (RT) for 24 h. Unreacted NHS-RHO was removed by membrane dialysis using Spectra/Por dialysis membrane (MWCO 3,500, Spectrum Laboratories Inc., Rancho Dominguez, CA) in excess deionized distilled water (ddH2O) for two days. The purified G4-RHO-NH2 conjugates were lyophilized over 2 days using a Labconco FreeZone 4.5 system (Kansas City, MO) and stored at −20 °C.
Next, FA was conjugated to G4-RHO-NH2 as described previously.4, 16 Briefly, FA (3.1 mg, 7.0×10−6 mol) was activated by EDC (13.4 mg, 7.0×10−5 mol) and NHS (8.0 mg, 7.0×10−5 mol) in 1.5 mL DMSO through vigorous stirring at RT for 1 h. The activated FA solution was added dropwise to 20 mg of G4-RHO-NH2 (1.4×10−6 mol) in 1 mL of ddH2O, followed by reaction under vigorous stirring at RT for 24 h. The product was purified by membrane dialysis as described above, resulting in G4-RHO-NH2.
The remaining primary amine groups of both G4-RHO-FA-NH2 and G4-RHO-NH2 were hydroxylated to minimize non-specific, electrostatic interactions with cell membranes,29–30 resulting in fully hydroxylated G4-RHO-FA-OH and G4-RHO-OH..
PEG-PLA was synthesized by ring opening polymerization of D,L-lactide as previously described.31 mPEG (100 mg and 150 mg) was transferred to a 3-neck round bottom flask and dried under vacuum for 2 h. D,L-lactide (1.0 g) and tin(II) 2-ethylhexanoate (30 mg) were added to the flask and dried under vacuum for an additional 1 h. The flask was placed in an oil bath that was pre-heated to 120 °C, and the polymerization was carried out under vigorous stirring for 4 h. The flask was then cooled to RT and 10 mL of DCM was added to dissolve the product. DCM was partially evaporated to adjust the solution viscosity, followed by precipitation in cold diethyl ether and vacuum drying overnight. The two feed ratios at 0.1:1 and 0.15:1 of mPEG:D,L-lactide resulted in PEG-PLA copolymers with MWs of 5K-45K and 5K-30K, respectively.
The hybrid NPs containing the various dendrimer conjugates were prepared using a double emulsion method as we previously described.28, 32 For example, G4-RHO-FA-OH (100 μL, 1 mg/mL in ddH2O) was added to 20 mg of either PEG5K-PLA23K, PEG5K-PLA30K, or PEG5K-PLA45K in 1 mL of DCM, and the mixture was sonicated for 1 min using a Misonix XL Ultrasonic Processor (100% duty cycle, 475 W, 1/8″ tip, QSonica, LLC, Newtown, CT). Two milliliters of 3% aqueous PVA solution was then added to the mixture, followed by additional sonication for 1 min. The double emulsion was poured into 20 mL of 0.3% PVA in ddH2O, and vigorously stirred at RT for 24 h to evaporate DCM. The resulting aqueous solution was transferred to Nalgene high-speed centrifuge tubes (Fisher Scientific, Pittsburg, PA) to remove PVA and unencapsulated G4-RHO-FA-OH by ultracentrifugation at 20,000 rpm for 30 min using a Beckman Avanti J25 Centrifuge (Beckman Coulter, Brea, CA). After washing the NPs five times with ddH2O, the pellet was resuspended in ddH2O, lyophilized over 2 days, and stored at −20 °C. All other dendrimer conjugates were encapsulated into PEG5K-PLA45K using the same method.
All dendrimer conjugates and PEG-PLA copolymers were characterized by 1H NMR using a 400 MHz Bruker DPX-400 spectrometer (Bruker BioSpin Corp., Billerica, MA) as described in our earlier publication.33 Particle size (diameter, nm) and surface charge (zeta potential, mV) of the conjugates and the nanohybrids were measured in triplicates by quasi-elastic laser light scattering using a Nicomp 380 Zeta Potential/Particle Sizer (Particle Sizing Systems, Santa Barbara, CA) in ddH2O. The measurements were performed using samples that were suspended in ddH2O at a concentration of 100 Tg/mL, filtered through a 0.45 μm syringe filter, and briefly vortexed prior to each measurement.
Loading was defined as the dendrimer conjugate content in the nanohybrids.28 Five milligrams of each nanohybrid formulation were completely dissolved in 1 mL of 0.5 M NaOH, followed by filtration through a 0.45 μm syringe filter. The fluorescence intensity from the filtrates was then measured using a SpectraMAX GeminiXS microplate spectrofluorometer (Molecular Devices, Sunnyvale, CA). The amount of the dendrimer conjugates in the filtrates was determined from a standard curve of each conjugate’s fluorescence versus concentration in 0.5 M NaOH. Loading was expressed as μg dendrimer conjugates per mg copolymer. Loading efficiency was defined as the ratio of the actual loading obtained to the theoretical loading.
Surface morphology of the nanohybrids was examined by scanning electron microscopy (SEM) using a JEOL-JSM 6320F field emission microscope (JEOL USA, Peabody, MA) as previously described.28 Samples were sputter-coated with Pt/Pd at a coating thickness of 6 nm (Polaron E5100 sputter coater system, Polaron, UK) and then visualized at an accelerating voltage of 4.0 mV and 8.0 mm working distance.
The release behaviors of G4-RHO-FA-OH from the prepared nanohybrids using different PEG-PLA copolymers (PEG5K-PLA23K (NP23), PEG5K-PLA30K (NP30), and PEG5K-PLA45K (NP45)) were studied in PBS and RPMI 1640.28 Five milligrams of each nanohybrid formulation were placed in microcentrifuge tubes and dispersed in 1 mL of either PBS (pH 7.4) or RPMI 1640 medium supplemented with 10% FBS in triplicates, and the solutions were placed in a shaking water bath (37 °C, 100 rpm). At various time points (1, 2, 4, 6, 8, 12, 24, 48 h; every other day thereafter), solutions were centrifuged at 10,000 rpm for 5 min, and the supernatants were collected. The nanohybrids were then redispersed in fresh PBS or RPMI 1640 medium and placed back in the water bath. The fluorescence of the supernatants was measured and the cumulative amount of G4-RHO-FA-OH released over time was determined from a standard curve of G4-RHO-FA-OH fluorescence versus concentration in PBS or RPMI 1640 medium.
The KB cell line was obtained from ATCC (Manassas, VA) and grown continuously as a monolayer in FA-deficient GIBCO RPMI 1640 medium (Invitrogen Corporation, Carlsbad, CA) to induce the overexpression of FR, under the same conditions that we previously reported.16 For confocal imaging, KB FR+ cells were seeded in 4-well chamber slides (Millicell EZ Slide, Millipore, Billerica, MA) at a density of 2.0×105 cells/well and incubated in FA-deficient RPMI 1640 for 24 h. The cells were treated with G4-RHO-FA-OH, G4-RHO-NH2, G4-RHO-OH, and the corresponding nanohybrids in the PEG5K-PLA45K shell (NP45) for 1 h and 4 h, at a concentration of 63 nM based on the dendrimer conjugates in PBS with Ca++ and Mg++ (Mediatech, Inc., Manassas, VA). Another group of KB FR+ was pre-incubated with 1 mM FA in PBS with Ca++ and Mg++ (from a stock solution of 100 mM FA in DMSO) for 30 min before adding G4-RHO-FA-OH and its NP45 formulation. Additionally, KB cells grown in complete RPMI 1640 (Invitrogen), resulting in FR-down-regulated KB (KB FR−), were used as a negative control16 and incubated with G4-RHO-FA-OH and its NP45 formulation. For the cellular interactions of the nanohybrids with different MWs of PLA, cells were similarly treated with G4-RHO-FA-OH and the corresponding nanohybrids prepared with PEG5K-PLA45K (NP45), PEG5K-PLA30K (NP30), and PEG5K-PLA23K (NP23) for 1, 4, and 8 h. After the treatments, cells were washed with PBS three times and fixed in 500 μL of 4% paraformaldehyde at RT for 10 min. The fixed cells were treated with antiphotobleaching mounting media with DAPI (Vector Laboratory Inc., Burlingame, CA), and covered with glass cover slips. Cellular binding and uptake were visualized using a Zeiss LSM 510 confocal laser scanning microscope (CLSM, Carl Zeiss, Germany). The 543 nm line of a 1 mW tunable argon laser was used for excitation of RHO, and a 25 mW diode UV 405 nm laser was used for excitation of DAPI. Emission was filtered at 565–595 nm and 420 nm for RHO and DAPI, respectively. Images were captured using a C-Apochromat 63×/1.2 W corr objective, with the pinhole set to 92 μm for the blue channel and 129 μm for the red channel. The detector gain was adjusted to 620 V for the blue channel and 824 V for the red channel. Images were acquired at a box size of 1024×1024 pixels, a scan speed of 7 fps (3.2 μs/pixel), and an average line scan of 16.
For the fluorescence activated cell sorter (FACS) analysis, KB FR+ were seeded in 12-well plates at a density of 1.0×106 cells/well and incubated in FA-deficient RPMI 1640 medium for 24 h. The cells were then treated under the same conditions described in the confocal observation above. After each incubation period, cells were washed with PBS and then suspended with trypsin/EDTA. Cell suspensions were centrifuged at 3500 rpm for 5 min, resuspended in 500 μL of 1% paraformaldehyde for fixation, and transferred to flow cytometry sample tubes. The fluorescence signal intensities from the samples were measured using a MoFlo cell sorter (BD, Franklin Lakes, NJ) and data analysis was performed using Summit v4.3 software (Dako Colorado, Fort Collins, CO).
In this study, we conducted a series of experiments to develop a nanohybrid system through encapsulation of targeted dendrimers into polymeric NPs. This novel platform was designed to achieve kinetically controlled receptor-specific interactions of FA-targeted dendrimers, which can be further modulated by varying MWs of PEG-PLA.
A general overview of the preparation of the hybrid NPs is illustrated in Figure 1. First, G4 PAMAM dendrimers were functionalized by sequential conjugation with RHO and FA, followed by hydroxylation of the remaining amine groups, resulting in G4-RHO-FA-OH (Figure 1(A)).4, 16 Note that the full hydroxylation step is critical to eliminate non-specific interactions between amine-terminated dendrimers and cells.7, 29, 34 The dendrimer conjugates were then encapsulated into PEG-PLA copolymers using the double emulsion method to produce nanohybrids (Figure 1(B)). Conjugation of RHO and FA to the dendrimers and successful end-capping of the amine groups was confirmed using 1H NMR and zeta potential measurements (Table 1 and Figure S1). The 1H NMR spectra (Figure S1) revealed that the conjugates prepared in this study contained approximately 3.9 and 4.3 RHO and FA molecules per dendrimer, respectively. Through various reactions, we prepared G4-RHO-NH2, G4-RHO-FA-OH, and the control conjugate G4-RHO-OH, to be hybridized with the PEG-PLA copolymers.
PEG-PLA copolymers were synthesized by bulk polymerization of lactide using mPEG5K as the initiator.31 1H NMR was used to confirm the chemical structure of the PEG-PLA copolymers and to estimate the MWs of the PLA block (Figure S2). By varying the mPEG:lactide feed ratio, two copolymers with different MWs of PLA were prepared. When the feed ratio of mPEG:lactide was 0.15:1, the MW of PLA was calculated to be 29,800 g/mol, based on the relative integration ratios of peak b around 3.62 ppm (the protons of the ethylene oxide repeating units) to peak c around 5.15 ppm (the lactide repeating units). This copolymer is referred to PEG5K-PLA30K throughout this paper. When the feed ratio of mPEG:lactide was 0.10:1, the MW of PLA was calculated to be 44,900 g/mol, which is referred to as PEG5K-PLA45K. A commercially available PEG5K-PLA23K was also used as a third copolymer for this study.
The first hypothesis was validated through the nanohybrid formation by double emulsion.28, 35 The encapsulation process yielded nanohybrids with controlled particle sizes around 100 nm in diameter and with high loading efficiencies (69 – 85%) (Table 1 and Figure 2). The significant differences in zeta potential for the prepared nanohybrids (−7.0 – −17.3 mV) from those of the dendrimer conjugates before encapsulation (3.4 – 28.1 mV) confirmed the successful encapsulation (Table 1). The results highlighted in Figure 2 and Table 1 support hypothesis (1) and clearly indicate that the encapsulation of the dendrimer conjugates into the polymeric NPs was successfully achieved.
For the second hypothesis, we tested if FR specificity of the FA-targeted dendrimer conjugates could be kinetically controlled by our nanohybrid design. The cellular interactions of the dendrimer conjugates and their respective nanohybrids were studied in KB FR+ cells using CLSM and FACS. Nanohybrids prepared with PEG5K-PLA45K (NP45) were employed for this experiment. Confocal images of KB FR+ cells after 1 h incubation (Figure 3) clearly show that only the targeted dendrimers (G4-RHO-FA-OH: red fluorescence) bind to the cell surface, which is in agreement with the FACS measurements (Figure 4) and a previous report.16 This interaction was completely blocked when the cells were pre-incubated with an excess amount of free FA (Figures 3 and and4),4), confirming that the observed dendrimer-cell interactions are a result of selective binding and uptake between FA on the dendrimers and FR on the cell surfaces.
In the case of the G4-RHO-FA-OH-based nanohybrids, we hypothesized that they will also bind specifically to the FR, but with a time delay as the targeted dendrimers are protected by the PEG-PLA NP shell. Indeed, receptor-specific binding was observed with the FA-targeted nanohybrids (G4-RHO-FA-OH/NP45) after 4 h incubation only. This interaction was comparable to that of the unencapsulated G4-RHO-FA-OH in terms of fluorescence intensity from the dendrimers in and/or on the cells. Furthermore, the dendrimer internalization was blocked upon pre-incubation with excess FA (Figures 3 and and4)4) and was negligible in KB FR− cells (Figure S3), confirming that the observed dendrimer interactions were FR-specific.
As shown in Figure 4, the control, non-targeted nanohybrids (G4-RHO-NH2/NP45 and G4-RHO-OH/NP45) showed a significantly lower, if not negligible, degree of cellular interaction. G4-RHO-NH2 exhibited a degree of non-specific interactions after 4 hr incubation likely due to electrostatic interactions.29, 34, 36 The significance of these results is two-fold: (1) FA-targeted dendrimers maintain selectivity to KB FR+ cells after the encapsulation and release process; and (2) the nanohybrid design allows for temporal control over the receptor targeting of the dendrimer-FA conjugates.
These observations highlight the potential of our nanohybrid system to control the selective cellular interaction kinetics of actively targeted polymer conjugates. This in turn would help address some of the challenges encountered with the currently available nanocarriers. We expect that by encapsulating the targeted dendrimers into a biodegradable polymeric shell, the conjugates would likely be shielded from immediate undesirable interactions with the receptors that are present in normal tissues. We also anticipate that through nanohybridization, the polymer conjugates would not be cleared from the circulation as rapidly as what has been previously observed with targeted polymers with surface-exposed FA.25 Current strategies that have been investigated to overcome the rapid clearance of dendrimer-drug conjugates include PEGylation to achieve steric stabilization, and surface modification by acetylation.18–19, 37 However, with proper optimization, the added advantage of our nanohybrid system is the possibility of exploiting two systems with different scales. That is, by encapsulating the dendrimer conjugates into polymeric NPs, the particle size of the system becomes large enough (~100 nm compared to 5–10 nm for the dendrimers) to potentially enable passive accumulation in the tumor tissues through the EPR effect. Furthermore, the targeted dendrimers upon release would likely penetrate into the solid tumors more effectively than the larger NPs, due to their small size.38–39 The in vivo validation of this nanohybrid design will be the subject of our future publications.
We tested the third hypothesis by encapsulating G4-RHO-FA-OH into PEG-PLA copolymers with various MWs of PLA, to produce three types of nanohybrids: G4-RHO-FA-OH/NP23, G4-RHO-FA-OH/NP30, and G4-RHO-FA-OH/NP45 prepared using PEG5K-PLA23K, PEG5K-PLA30K, and PEG5K-PLA45K, respectively. The release profiles of the dendrimer conjugates in PBS buffer and in serum-containing RPMI 1640 (Figure 5) were biphasic, consisting of a relatively fast release within the first 8 h, most probably due to surface desorption of the dendrimer conjugates.40 This early release can also indicate that a portion of the conjugates may not have been completely encapsulated within the NP core. The profile became slow and sustained afterward, likely resulting from the gradual degradation of the copolymer matrix and diffusion of the dendrimer conjugates.41 The MW of the PLA block was inversely proportional to the release rate of the dendrimer conjugates regardless of the release medium, which is similar to previous reports.42 The higher PEG:PLA ratio in the smaller MW PLA copolymers also results in increased hydrophilicity, greater water uptake, and faster degradation compared to copolymers with a smaller PEG:PLA ratio.31, 43 The smaller MW and the more hydrophilic surface are expected to contribute to the faster release profile of the G4-RHO-FA-OH/NP23 nanohybrids than that of the NP30 and NP45 nanohybrids. The degradation kinetics of PEG-PLA copolymers themselves are reportedly slower than the release kinetics observed in this study.44–45 This suggests that the release of the dendrimer conjugates is not only governed by degradation of the PLA block, but by diffusion as well, similar to what was observed for proteins with comparable MW.46–49
We observed much faster release profiles in serum-containing RPMI 1640 compared to PBS, (Figure 5(B)), likely attributed to the presence of serum proteins such as albumin in the RPMI 1640. Serum albumin has been shown to exhibit esterase-like activity that accelerates the hydrolysis of the PLA chains.43, 50–51 Regardless of the release medium, the overall release profiles demonstrate that the release kinetics of the targeted dendrimers can be controlled by varying the MW of the encapsulating copolymers, which would in turn affect the cellular targeting and interaction kinetics.
In order to achieve kinetic control over FR targeting, each of the G4-RHO-FA-OH/NP23, G4-RHO-FA-OH/NP30, and G4-RHO-FA-OH/NP45 nanohybrids was incubated with KB FR+ cells up to 8 h. After 1 h incubation (Figures 6 and and7),7), only the G4-RHO-FA-OH/NP23 nanohybrids showed noticeable receptor interactions, which were comparable to unencapsulated G4-RHO-FA-OH. NP30 and NP45 nanohybrids exhibited slower cellular interaction and targeting kinetics. After 4 h incubation, they started to show significant receptor binding, but these interactions were still lower than those observed with the NP23 nanohybrids. After 8 h of incubation, the FACS results demonstrate that all nanohybrids show similar degrees of cellular binding/uptake to unencapsulated G4-RHO-FA-OH (Figure 7). The observed temporal delay in the cellular interactions of the NP30 and NP45 nanohybrids can be attributed to the increase in the MW of the PLA block, which delays the release of the FA-conjugates to interact with FR on the cell surface. The larger differences in the release kinetics (Figure 5) compared to the differences in cellular interaction kinetics (Figures 6 and and7)7) could be due to the rapid desorption of G4-RHO-FA-OH located near the nanohybrid surfaces and the facilitated degradation in the presence of KB cells.52 Overall, it is obvious that the kinetics of the cellular interactions of the nanohybrids are primarily governed by the controlled release profiles of the different MW PLAs, further supporting our third hypothesis.
Taken together, the results from this study support the three hypotheses stated earlier. For hypothesis (1), both targeted and non-targeted dendrimer conjugates were encapsulated into PEG-PLA NPs with controlled sizes at high encapsulation efficiencies (Table 1 and Figure 2). Hypothesis (2) was validated by the FR specificity and temporally controlled targeting of the FA-targeted dendrimer-containing nanohybrids (Figures 3 and and4).4). The release tests and the cellular interaction studies using CLSM and FACS (Figures 5, ,6,6, and and7)7) demonstrated that the targeting kinetics of the dendrimers can be controlled by altering the MW of the biodegradable components of the nanohybrid systems, thereby supporting hypothesis (3). This control over the targeting kinetics further emphasizes the versatility and flexibility of our nanohybrid design, which can be tailored to achieve the desired targeting kinetics and release rates of multifunctionalized polymer-drug conjugates. Our novel nanohybrids demonstrate a great potential to enable precise control over the targeting kinetics of the nanocarrier to tumor cells, which can potentially minimize premature elimination and off-target delivery of the conventional nanocarriers that have targeting agents being exposed on the surface. Additionally, the flexibility of the nanohybrid design can also be exploited to achieve temporally controlled ligand presentations for targeting inflammatory diseases.53 The staged, temporally controlled targeting to multiple ligands (e.g. selectins, IL-6, STAT3) would effectively block the inflammatory responses that frequently accompany various cancers.54–55
This work was partially supported by Vahlteich research funds awarded to SH from the University of Illinois College of Pharmacy. This work was conducted in a facility constructed with support from grant C06RR15482 from the NCRR, NIH.