Nanoparticles coated with high molecular weight PEG penetrate mucus and provide uniform vaginal and colorectal distribution in vivo

Abstract

Aim:

We previously reported that nanoparticles (NPs) coated with 10 kDa PEG were mucoadhesive. Here, we demonstrate that by increasing the surface density, PEG with molecular weight (MW) as high as 40 kDa can be used as a mucoinert NP surface coating.

Materials & methods:

We compared two sets of reaction conditions for coating model polystyrene NPs with 10 kDa PEG and used optimized conditions to coat NPs with PEG as high as 40 kDa in MW. We then characterized NP transport in human cervicovaginal mucus ex vivo. We further administered PEG-coated NPs to the mouse cervicovaginal tract and colorectum to assess mucosal distribution in vivo.

Results & conclusion:

We demonstrate here that PEG with MW as high as 40 kDa can be densely grafted to the surface of NP to prevent interactions with mucus. NP coated with 10–40 kDa PEG rapidly diffused through human cervicovaginal mucus ex vivo, and uniformly lined the mouse colorectal and vaginal epithelium in vivo.

Background

Mucus, the first-line defense covering all mucosal surfaces, is an adhesive, viscoelastic gel that effectively traps many pathogens, thereby preventing them from reaching and infecting the underlying cells and tissues. Mucus also prevents particulates from reaching the epithelial surface if they are larger than, and/or adhere to, the mucus mesh [1–6], which can decrease drug delivery efficiency. Many nanoparticle (NP) formulations are designed to be mucoadhesive, but we have shown that adhesion to mucus leads to limited distribution over vaginal, lung, and colorectal tissues [6–9]. In contrast, NPs densely coated with nonmucoadhesive hydrophilic polymers, such as PEG, are able to penetrate mucus barriers to reach and uniformly coat epithelial surfaces [4,6,8,9]. In addition, these mucus penetrating nanoparticles (MPPs) are retained for longer periods of time in the cervicovaginal and respiratory tracts compared with mucoadhesive particulates [6,8], indicating that MPPs may be more suitable for distributing drugs to the entire epithelial surface and providing prolonged drug retention [6,8,9].

Prior work demonstrated that incorporating PEG onto the surfaces of hydrogels and hydrogel microparticles encouraged adhesion to mucus [10–16]. PEG was described to act as a mucoadhesive ‘glue’ that could interpenetrate and entangle with mucin fibers [11,13] or form hydrogen bonds to the carbohydrate regions of the mucin fibers [17]. In our previous work, we hypothesized that the hydrophilic and uncharged nature of PEG could also be used to prevent mucoadhesion, but only if low-molecular-weight (MW) PEG chains were grafted onto the surface of NPs at high density to prevent entanglement and interaction with mucins [4]. Indeed, we showed that carboxylate-modified polystyrene (PS) NPs densely coated with 2 or 5 kDa PEG (PS-PEG) rapidly diffused in human cervicovaginal mucus (CVM) [4,18–20], whereas NPs with positively charged surfaces, and/or uncoated hydrophobic surfaces were mucoadhesive [4,6]. When we coated PS with 10 kDa PEG, the PS-PEG NPs were immobilized in human CVM, which we hypothesized was due to the increased PEG MW and/or insufficient PEG coverage on the NP surface [21]. To explore the latter issue, here we optimized the PEGylation reaction conditions to achieve increased PEG grafting densities on the surface of PS NPs compared with our previous methods [1,22]. We found that PS NPs can be densely coated with PEG with MW as high as 40 kDa, leading to rapid diffusion in human CVM ex vivo and uniform mucosal distribution in the mouse vagina and colorectum in vivo following topical administration. The results described here redefine and broaden the design criteria for NP systems for improved mucosal delivery.

Experimental section

NPs were prepared using methods previously described [1,4]. Briefly, 40, 100 and 200 nm carboxylate-modified PS beads were coated with 5 kDa up to 40 kDa methoxy-PEG-amine using buffers at pH 6 and 7.4. Particle size and surface ζ-potential were analyzed using dynamic light scattering and laser Doppler anemometry. Multiple particle tracking experiments were performed using freshly obtained, undiluted human CVM and epithelial distribution images were obtained by administering NPs in hypotonic solutions to the vagina of naturally cycling mice in estrus phase or to the colorectum of mice that had been starved for 24 h [6,9,23]. All animal experiments were approved by the Johns Hopkins University Animal Care and Use Committee. Further details can be found in the Supplementary materials.

Results & discussion

We first directly compared our prior method (2-(N-morpholino)ethanesulfonic acid, or MES, buffer method) [21] and the recently optimized method (borate buffer method) [1] for PEG coating density achieved on PS NPs. Similar to our previous observations, PS NPs coated with 10 kDa PEG using the MES method (PS-PEG_MES, 220 ± 10 nm; -12 ± 0.4 mV) were adhesively immobilized in CVM; the ensemble averaged mean-squared displacement (<MSD>) for PS-PEG_MES NP was >10,000-fold slower in CVM compared with the theoretical diffusion rates of the similaly sized NPs in water (Table 1). In contrast, we found that the same PS NPs coated with 10 kDa PEG using the borate method (PS-PEG_borate, 260 ± 7 nm; -0.7 ± 0.5 mV) rapidly diffused in CVM. The ensemble-averaged mean-squared displacement, <MSD>, of PS-PEG_borate NPs was only 11-fold slower in CVM compared with their theoretical diffusion rate in water (Table 1), and approximately 1000-fold higher than either uncoated PS NPs (180 ± 1 nm; -59 ± 2 mV) or PS-PEG_MES (Figure 1A & Table 1). The <MSD> (Figure 1A) and the logarithmic distribution of individual MSD values (Figure 1B) for PS-PEG_MES NP were similar to uncoated PS NPs, as we reported previously [21]. NP trajectories representing 3 s of movement in CVM further emphasize the difference in transport behavior, as the motion of PS-PEG_borate NPs reflected diffusive motion, while the trajectories of PS and PS-PEG_MES NPs were highly constrained (Figure 1C).

Table 1. . Size, ζ-potential, PEG surface density (area covered by PEG/total surface area, or Γ/SA), and the comparison of the ensemble averaged mean-squared displacement in mucus (<MSD>) to the theoretical MSD of similarly sized particles in water (MSD_w) (indicating how much slower the nanoparticles move in mucus, MSDw/<MSD>), of 100 and 200 nm polystyrene and PS-PEG NPs prepared by various methods.

Size (nm)	Type	PEG MW (kDa)	Hydrodynamic diameter (nm)	ζ-potential (mV)	Γ/SA	# PEG chains/100 nm2	MSDw/<MSD>
40	PS	-	56 ± 2	-33 ± 0.6	NA	NA	NA
	PS-PEG	5	60 ± 1	-2.2 ± 0.2	NA	NA	NA
		10	68 ± 0.4	-2.9 ± 0.4	NA	NA	NA
		20	84 ± 3	-2.6 ± 0.2	NA	NA	NA
		40	97 ± 1	-3.4 ± 0.5	NA	NA	NA
100	PS	-	90 ± 1	-51 ± 1.6	0	0	>10,000
	PS-PEG	5	110 ± 2	-3.1 ± 0.3	>2†	˜9†	10
		10	120 ± 7	-0.5 ± 0.1	2.0 ± 0.1	4.4 ± 0.2	18
		20	130 ± 4	-0.4 ± 0.1	3 ± 0.1	3.3 ± 0.1	6
		40	170 ± 8	-1 ± 0.1	2.1 ± 0.2	1.2 ± 0.1	20
200	PS	-	180 ± 1	-59 ± 2	0	0	>10,000
	PS-PEG	5	230 ± 5	-1.6 ± 0.1	1.6 ± 0.1	7.1 ± 0.4	9
		10	260 ± 7	-0.7 ± 0.5	1.5 ± 0.1	3.3 ± 0.1	11
		10 (MES)	220 ± 10	-12 ± 0.4	1.3 ± 0.1	2.9 ± 0.3	>10,000
		20	270 ± 7	-2 ± 0.7	1.7 ± 0.1	1.9 ± 0.2	10
		40	300 ± 5	-1 ± 0.6	1.7 ± 0.1	1 ± 0.1	15

^†Based on literature [1].

Unless otherwise indicated, PS-PEG NPs were prepared via the borate method. Values are averaged over n ≥ 3 samples. MSD: Mean-squared displacement; MW: Molecular weight; NP: Nanoparticle; PS: Polystyrene.

Figure 1. . Transport of 200 nm polystyrene and PS-PEG nanoparticles coated with 10 kDa PEG using the borate or MES method.

(A) Ensemble mean-squared displacement (<MSD>) with respect to time up to 3 s for PS and PS-PEG NPs, including the theoretical MSD of 200 nm NPs in water (W). (B) Distributions of the logarithms of individual particle MSD of PS and PS-PEG NPs at a time scale of 1 s. (C) Representative trajectories for 3 s of motion of PS and PS-PEG NPs. Data are representative of n ≥ 3 samples.

NP: Nanoparticle; PS: Polystyrene.

We next sought to quantify the PEG surface density on the PS-PEG NPs. We previously used an indirect method involving conjugation of fluorescent dyes to the unreacted carboxylic acid groups remaining on the NP surface after the PEG conjugation [21]. Using this method, we approximated that approximately 69% of carboxylic acid groups on the PS NP surface were conjugated to 2 kDa PEG, and approximately 65% of the carboxylic acid groups on the PS NP surface were conjugated to 10 kDa PEG [21]. Here, we used a more quantitative nuclear magnetic resonance method that can also be used to infer the physical conformation of the PEG chains based on the packing density (Γ/SA, where Γ is the unconstrained surface area that would be covered by the grafted PEG chains and SA is the total NP surface area) [1,22]. We calculated an increased Γ/SA for the PS-PEG_borate NPs (1.5 ± 0.1), compared with 1.3 ± 0.1 for PS-PEG_MES NPs (Table 1), indicating that the PEG chains were more densely packed on the surfaces of the PS-PEG_borate NPs. The Γ/SA values correspond to 3.3 ± 0.1 10 kDa PEG chains per 100 nm² for PS-PEG_borate NPs, compared with 2.9 ± 0.3 10 kDa PEG chains per 100 nm² for PS-PEG_MES NPs. Our results imply that there is a narrow margin between where an NP surface coating with PEG is sufficient or insufficient for precluding interactions with mucus. We similarly found in prior work that a Γ/SA of at least 2.0 was required for 100 nm PS NPs coated with 5 kDa PEG to effectively penetrate through the brain extracellular matrix [1], whereas Γ/SA = 1.7 was an insufficient coating, implying that the required density of the PEG coating depends on the surrounding environment, the NP size and the PEG MW, among other factors. Furthermore, we recently observed that biodegradable NPs composed of block copolymers of PEG (MW as high as 10 kDa) and poly(lactic-co-glycolic acid) were sufficiently densely coated with PEG to allow rapid diffusion in human CVM [22]. Since the poly(lactic-co-glycolic acid) NPs were formed using an emulsion method that allowed PEG to partition to the NP surface during the slow hardening process, rather than grafting PEG onto the surface of preformed NPs, Γ/SA values of 2.3 and greater (Γ/SA = 3.0 for 10 kDa PEG) were achieved [22].

Using the optimized borate method, we then explored the issue of whether an upper limit to the PEG MW could be found where PEG-coated PS NPs would become mucoadhesive. We found that 200 nm PS NPs coated with 5–40 kDa PEG using the borate method (PS-PEG_5–40kDa, final sizes 230–300 nm; Table 1) were able to rapidly penetrate CVM, as indicated by the high <MSD> values measured (Figure 2A, 200 nm). In addition, the 3 s trajectories for all PS-PEG_5–40kDa NPs reflected diffusive behavior, in stark contrast to the uncoated PS NPs (Figure 2B; 200 nm). The PS-PEG_5–40kDa NPs also had high PEG density, Γ/SA ≥1.5 (Table 1). It was evident that the number of PEG molecules per 100 nm² decreased as the PEG MW increased, which can be attributed to the increased amount of space occupied by each PEG chain as the MW increases; the area occupied by one unconstrained 5 kDa PEG chain is approximately 23 nm², compared with approximately 180 nm² for unconstrained 40 kDa PEG molecules. Additionally, the hydrodynamic diameter of the NPs after PEGylation increased as the PEG MW increased; dense packing of the PEG chains causes elongation, so higher MW PEG chains would create a thicker corona.

Figure 2. . Transport of 100 and 200 nm polystyrene and PS-PEG nanoparticles coated with varying molecular weight PEG in cervicovaginal mucus.

(A) Ensemble averaged mean-squared displacement (<MSD>) as a function of time up to 3 s for PS and PS-PEG NPs coated with 5, 10, 20 and 40 kDa PEG using the borate method, including the theoretical MSD of 100 or 200 nm particles in water (W). (B) Representative trajectories for 3 s of motion of PS and PS-PEG NPs in CVM. Data are representative of n ≥ 3 samples.

CVM: Cervicovaginal mucus; NP: Nanoparticles; PS: Polystyrene.

We then investigated the role of the PS NP core size. We coated 100 nm PS NPs with 5–40 kDa PEG using the borate method (PS-PEG_5–40kDa, final sizes 110–170 nm, Table 1). As shown in Table 1, the Γ/SA values were all >2 for 100 nm PS-PEG_5–40kDa NP, resulting in rapid NP diffusion through CVM (Figure 2A, 100 nm). PEG-coated NP trajectories reflected diffusive motion, in stark contrast to the uncoated PS NPs (Figure 2B, 100 nm). All 100 and 200 nm PS-PEG_5–40kDa NP formulations were slowed <20-fold in CVM compared with their theoretical diffusion rates in water (Table 1).

As previously described, PEG can also be used at lower grafting densities to enhance mucoadhesion [11–14,16,17,24,25]. Peppas et al. have suggested that tethering PEG to different polymers and gels makes them more mucoadhesive by interpenetrating into the mucus gel [11,13,17]. DeAscentiis et al. determined that for poly(hydroxyethyl methylcellulose) microparticles, 1 kDa PEG coatings led to increased mucoadhesion compared with 100 kDa PEG [10]. They hypothesized that because 100 kDa PEG chains could form more self-entanglements than 1 kDa PEG chains, there would be less interpenetration of 100 kDa PEG chains into the mucin gel, and thus, less mucoadhesion [10]. In our case, we hypothesize that by densely packing the PEG chains on the NP surface, the PEG becomes aligned and sterically restricted from penetrating into the mucin gel, regardless of PEG MW. Adhesion has been shown to depend on PEG chain movement, and thus if chains are constrained enough, they will not be able to interpenetrate with the mucus mesh, leading to decreased mucoadhesion [17].

Similarly, it is well known that dense PEG coatings increase the circulation time of systemically administered NPs by preventing protein adsorption to the NP surface and interactions with cells [26–29]. Yang et al. found that <20% of NPs coated with up to 20 kDa PEG were cleared from systemic circulation for Γ/SA >2.8. At this PEG surface density, they determined that neighboring PEG chains were highly unlikely to simultaneously reach an extended configuration, which could expose the particle surface to protein adsorption. In addition, they found that even a minor decrease in PEG density led to a decrease in systemic circulation time to <2 h [29]. Similarly, we found that a small decrease in NP PEG surface density resulted in a transition from nonmucoadhesive surface properties to mucoadhesive surface properties. The exact transition point between adhesive and nonadhesive is likely to depend on the surrounding environment, the NP size and core material, and the PEG MW.

Lastly, we explored the impact of PEG MW on the distribution of PS-PEG NPs in vivo. We have previously observed that the diffusion of NPs in mucus ex vivo correlates well with the observed distribution at mucosal surfaces in vivo. In other words, particles that diffused rapidly in mucus ex vivo also distributed more uniformly and persisted longer upon administration locally to a mucosal surface. Further, we have observed that improved NP distribution provides more efficacious treatment and prevention of diseases at mucosal sites, including prevention of herpes (HSV-2) infection in the cervicovaginal tract, prevention of asthma-induced lung inflammation, and treatment of cervical cancer [6,8,30,31]. Here, we found that PS-PEG NPs densely coated with as high as 40 kDa PEG also distributed uniformly in the cervicovaginal tract and colorectum of mice (Figure 3). We administered PS-PEG NPs of appropriate sizes to distribute uniformly in the mouse vagina (˜100 nm) and colorectum (˜40 nm) when they are sufficiently well PEGylated and administered in a hypotonic vehicle that induces fluid absorption by the epithelium [6,7,9,23]. Similar to our prior observations, uncoated PS NPs aggregated in the luminal mucus layers, but all PS-PEG NP formulations with sufficiently dense PEG coatings for diffusion in ex vivo mucus samples (PS-PEG_5–40kDa) were transported rapidly and uniformly to the vaginal and colorectal epithelial surfaces in vivo (Figure 3). In addition to further highlighting the importance of PEG surface density and addressing the mechanisms involved in the PEG mucoadhesivity paradox, our results suggest that up to 40 kDa MW PEG may in fact be used to make MPP formulations.

Figure 3. . Distribution of polystyrene and PS-PEG nanoparticles coated with varying MW PEG in the mouse vagina and colorectum.

Transverse vaginal and colorectal tissue cryosections obtained 5–10 min after administration of solutions containing either PS, or PS-PEG NPs coated with various MW PEG (5, 10, 20 and 40 kDa). Cell nuclei are stained blue with DAPI and scale bars represent 300 μm. Images are representative of n ≥ 3 mice. MW: Molecular weight; NP: Nanoparticle; PS: Polystyrene.

Executive summary.

Background

• Mucus is the first-line defense that protects the epithelium.

• Mucus effectively traps particulates and pathogens, including nanoparticles (NPs).

• PEG has been shown to prevent NPs from adhering to mucus when grafted densely onto the particle surface.

• Previous work indicated that only lower molecular weights (MWs) of PEG (2–5 kDa) could prevent mucoadhesion.

Results & discussion

• Using improved reaction conditions, NPs were sufficiently densely coated with PEG up to 40 kDa MW, allowing coated NPs to rapidly diffuse through human cervicovaginal mucus.

• More effective mucus penetration by NPs more densely coated with PEG with MW up to 40 kDa indicates that the physical conformation of PEG on the NP surface is key for preventing interactions with mucus.

• Densely PEG-coated nanoparticles, regardless of PEG MW, uniformly lined the mouse colorectal and vaginal epithelium in vivo.

Conclusion

• PEG MW of up to 40 kDa can be used to formulate mucus penetrating nanoparticles.

• This work broadens the design criteria for designing drug and gene delivery systems for improved mucosal delivery.