The surface of a polymeric nanoparticle (NP) is often functionalized with cell-interactive ligands and/or additional polymeric layers to control NP interaction with cells and proteins. However, such modification is not always straightforward when the surface is not chemically reactive. For this reason, most NP functionalization processes employ reactive linkers or coupling agents or involve pre-functionalization of the polymer, which are complicated and inefficient. Moreover, pre-functionalized polymers can lose the ability to encapsulate and retain a drug if the added ligands change chemical properties of the polymer. To overcome this challenge, we use dopamine polymerization as a way of functionalizing NP surfaces. This method includes brief incubation of the pre-formed NPs in a weak alkaline solution of dopamine, followed by secondary incubation with desired ligands. Using this method, we have functionalized poly(lactic-co-glycolic acid) (PLGA) NPs with three representative surface modifiers: a small molecule (folate), a peptide (Arg-Gly-Asp), and a polymer [poly(carboxybetaine methacrylate)]. We confirmed that the modified NPs showed the expected cellular interactions with no cytotoxicity or residual bioactivity of dopamine. The dopamine polymerization method is a simple and versatile surface modification method, applicable to a variety of NP drug carriers irrespective of their chemical reactivity and the types of ligands.

Stringing them along: The pores of a mesoporous silica particle were filled with guest molecules and then blocked by threading cyclodextrin molecules (CDs) onto the surface‐grafted polyethylenimine (PEI) chains at pH 11. At pH 5.5, the guest molecules can be released from the pores of the particle by reversible dethreading of the CDs from the PEI chains.

We report biocompatible, cell-permeable core-shell-corona polymer micelles bearing glutathione-cleavable shell cross-links, which allow the facilitated release of entrapped anticancer drugs at cytoplasm in response to an intracellular glutathione level.

We report a theranostic nanoparticle that can express ultrasound (US) imaging and simultaneous therapeutic functions for cancer treatment. We developed doxorubicin-loaded calcium carbonate (CaCO3) hybrid nanoparticles (DOX-CaCO3-MNPs) through a block copolymer templated in situ mineralization approach. The nanoparticles exhibited strong echogenic signals at tumoral acid pH by producing carbon dioxide (CO2) bubbles and showed excellent echo persistence. In vivo results demonstrated that the DOX-CaCO3-MNPs generated CO2 bubbles at tumor tissues sufficient for echogenic reflectivity under a US field. In contrast, the DOX-CaCO3-MNPs located in the liver or tumor-free subcutaneous area did not generate the CO2 bubbles necessary for US contrast. The DOX-CaCO3-MNPs could also trigger the DOX release simultaneously with CO2 bubble generation at the acidic tumoral environment. The DOX-CaCO3-MNPs displayed effective antitumor therapeutic activity in tumor-bearing mice. The concept described in this work may serve as a useful guide for development of various theranostic nanoparticles for US imaging and therapy of various cancers.

Amphiphilic diblock copolymers were synthesized based on poly(2-ethyl-2-oxazoline) (PEtOz) as a hydrophilic block and aliphatic polyesters such as poly(l-lactide) (PLA) or poly(ε-caprolactone) (PCL) as a hydrophobic block. Their micellar characteristics in an aqueous phase were investigated by using dynamic light scattering and fluorescence techniques. The block copolymers formed micelles in the aqueous phase with critical micelle concentrations (cmcs) in the range of 1.0−8.1 mg/L. The cmc values become lower upon increasing the length of the hydrophobic block. The mean diameters of the micelles were in the range of 108−192 nm, with a narrow distribution. In general, the micelle size increased as the hydrophobic PLA or PCL block became larger. The partition equilibrium constants, K v, of pyrene in the micellar solutions of the block copolymers were from 1.79 × 105 to 5.88 × 105. For each block copolymer system of PEtOz−PLA or PEtOz−PCL, the K v value increased as the length of the hydrophobic block increased. The steady-state fluorescence anisotropy values (r) of 1,6-diphenyl-1,3,5-hexatriene (DPH) were 0.265−0.284 in PEtOz−PLA solution and 0.189−0.196 in PEtOz−PCL solution. The anisotropy values of PEtOz−PLAs were higher than those of PEtOz−PCLs. The anisotropy values were independent of the length of the hydrophobic block when the chemical structures of the hydrophobic blocks were identical. The micelles underwent hydrogen bonding at pH <3.5 with poly(acrylic acid), which produced polymer complex precipitates that could be reversibly dispersed as micelles at pH >3.8.

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Since 1960 when the history of modern hydrogels began significant progresses have been made in the field of controlled drug delivery. In particular, recent advances in the so-called smart hydrogels have made it possible to design highly sophisticated formulations, e.g., self-regulated drug delivery systems. Despite intensive efforts, clinical applications of smart hydrogels have been limited. Smart hydrogels need to be even smarter to execute functions necessary for achieving desired clinical functions. It is necessary to develop novel hydrogels that meet the requirements of the intended, specific applications, rather than finding applications of newly developed hydrogels. Furthermore, developing smarter hydrogels that can mimic natural systems is necessary, but the fundamental differences between natural and synthetic systems need to be understood. Such understanding will allow us to develop novel hydrogels with new, multiple functions we are looking for.