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Keywords:

  • polypeptides;
  • stimuli-sensitive materials;
  • drug delivery;
  • gene delivery;
  • poly(amino acids)

Abstract

  1. Top of page
  2. Abstract
  3. 1. Introduction
  4. 2. Synthesis of Polypeptides
  5. 3. Stimuli-Sensitive Polypeptide-Based Materials
  6. 4. Stimuli-Sensitive Functionalized Polypeptides
  7. 5. Conclusions and Perspectives
  8. Acknowledgements
  9. Biographical Information

Stimuli-sensitive synthetic polypeptides are unique biodegradable and biocompatible synthetic polymers with structures mimicking natural proteins. These polymers exhibit reversible secondary conformation transitions and/or hydrophilic–hydrophobic transitions in response to changes in environmental conditions such as pH and temperature. The stimuli-triggered conformation and/or phase transitions lead to unique self-assembly behaviors, making these materials interesting for controlled drug and gene delivery applications. Therefore, stimuli-sensitive synthetic polypeptide-based materials have been extensively investigatid in recent years. Various polypeptide-based materials, including micelles, vesicles, nanogels, and hydrogels, have been developed and tested for drug- and gene-delivery applications. In addition, the presence of reactive side groups in some polypeptides facilitates the incorporation of various functional moieties to the polypeptides. This Review focuses on recent advances in stimuli-sensitive polypeptide-based materials that have been designed and evaluated for drug and gene delivery applications. In addition, recent developments in the preparation of stimuli-sensitive functionalized polypeptides are discussed.


1. Introduction

  1. Top of page
  2. Abstract
  3. 1. Introduction
  4. 2. Synthesis of Polypeptides
  5. 3. Stimuli-Sensitive Polypeptide-Based Materials
  6. 4. Stimuli-Sensitive Functionalized Polypeptides
  7. 5. Conclusions and Perspectives
  8. Acknowledgements
  9. Biographical Information

Controlled drug release systems are designed to deliver therapeutical drugs to desirable sites at appropriate time.1–3 In recent years, stimuli-sensitive polymers that respond to changes in environmental conditions,3, 4 such as pH, temperature, salt, light, biomolecules, and electromagnetic field, have received extensive attention for their unique advantages in applications for controlled drug delivery. The release of drugs from the smart delivery systems can be triggered by the specific environments of some organs, intracellular space, or pathological sites, leading to an enhanced specificity of drug delivery and less side effects.5, 6 In addition to intelligent drug-release behaviors, drug carriers with good biocompatibility and appropriate biodegradability are preferred in practical applications.7

Synthetic polypeptides, which are poly(amino acid)s linked by peptide bonds, are unique biodegradable and biocompatible synthetic polymers with structures mimicking natural proteins.8 In comparison with conventional biodegradable polymers, synthetic polypeptides may form stable secondary structures, such as α-helix and β-sheet, due to cooperative hydrogen-bonding, leading to unique self-assembly behaviors.9, 10 In addition, the self-assembly structures of some polypeptides exhibit transitions in response to external stimuli, such as pH, salt, and temperature. Especially, some polypeptides with ionizable side groups can form electrostatic interactions with oppositely charged drugs and bioactive macromolecules, such as DNAs, RNAs, and proteins, and functional moieties can be incorporated into the side chains of the polypeptides. Therefore, polypeptide-based materials have attracted increasing attention for their great potential in biomedical and pharmaceutical applications.11–14 It is worth mentioning that polymers that exhibit sharp phase and structure transitions in response to physiologically relevant stimuli have advantages in drug-delivery applications. Typical physiologically relevant stimuli include the difference in microenvironment between normal tissues and some pathological sites, e.g., tumor sites,5, 15 the difference in microenvironment, such as pH and redox environment,5, 15–17 between the extracellular and intracellular space, as well as the pH shift between the stomach (pH ∼ 2.0) and the intestine (pH ∼ 7.0).6 In addition, polymers that exhibit temperature-dependent phase transitions at around physiological temperature, such as 10–42 °C,18, 19 are also interesting for controlled drug delivery. Consequently, polypeptide-based materials capable of responding to physiologically relevant stimuli, including pH, temperature, redox environment, and dual stimuli, have been extensively developed for drug delivery applications in recent years, and many promising results have been reported. Based on the polypeptides with different architectures and compositions, intelligent polymeric materials, including micelles, vesicles, nanogels, and hydrogels, have been developed, as schematically illustrated in Figure1. This Review focuses on recent progress in stimuli-responsive polypeptide-based materials that have been designed and evaluated for drug- and gene-delivery applications, including polypeptide-based micelles, vesicles, nanogels, and hydrogels. In addition, recent developments in the preparation of stimuli-sensitive functionalized polypeptides are discussed.

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Figure 1. Schematic illustration of intelligent polypeptide-based materials, including micelles, vesicles, nanogels, and hydrogels, as well as functionalized polypeptides.

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2. Synthesis of Polypeptides

  1. Top of page
  2. Abstract
  3. 1. Introduction
  4. 2. Synthesis of Polypeptides
  5. 3. Stimuli-Sensitive Polypeptide-Based Materials
  6. 4. Stimuli-Sensitive Functionalized Polypeptides
  7. 5. Conclusions and Perspectives
  8. Acknowledgements
  9. Biographical Information

Polypeptides are mainly prepared via ring opening polymerization (ROP) of α-amino acid N-carboxyanhydrides (NCAs) by using amine-based initiators.8, 20, 21 The typical mechanisms of amine-initiated ROP of NCAs include “amine mechanism (AM)” and “activated monomer mechanism (AMM)”.20, 21 The ROP of NCAs via AM is believed to be performed by attacking CO-5 of the NCAs using primary amines or some highly nucleophilic secondary amines (Scheme1a).20 On the other hand, the ROP via “activated monomer mechanism” is achieved by deprotonation of NH-3 of the NCAs using tertiary amines or some secondary amines with bulkyl substitution groups (Scheme 1b).20, 21 The ROP of NCAs initiated by primary amines leads to the successful preparation of different kinds of polypeptides and polypeptide-based copolymers with desired molecular weights (MW) and architectures, while that initiated via AMM usually results in the formation of polypeptides with high MW.

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Scheme 1. The ring opening polymerization of NCAs via “amine mechanism” (a) and “activated monomer mechanism” (b).20, 21

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Even though some side reactions, e.g., end-group cyclization, may occur during the primary amine-initiate-polymerization, causing contamination of polypeptide products, considerable approaches have been developed recently to prepare polypeptides with controlled MW and polydispersities (PDIs) as well as well-defined topological structures. Deming and co-workers have used a series of transition metal complexes as initiators to obtain well-defined polypeptides and block copolypeptides.8 Hadjichristidis and co-workers have obtained well-defined polypeptides using conventional amines as initiators under high vacuum conditions.22 Schlaad et al. and Lutz et al. have reported that ammonium chloride-functionalized macro-initiators can effectively suppress side reactions,23, 24 leading to the formation of well-defined polypeptide-based block copolymers. In addition, controlled ROP of NCAs has been achieved by using organosilicon.25

In the past decade, amphiphilic polypeptide-based block copolymers have been extensively studied for biomedical applications, attributed to the fact that they can self-assemble into different stimuli-sensitive nanometer-sized or micrometer-sized structures in aqueous solution and the size and morphology of these materials can be readily tuned by varying the composition of the block copolymers.26–28 Accordingly, different amphiphilic polypeptide-based block copolymers have been synthesized for biomedical applications. The most widely used method to synthesize polypeptide-based block copolymers is the ROP of NCAs by using polymeric macro-initiators. In addition, block copolymers are also prepared by highly effective coupling reactions, such as click chemistry and carbondiimide chemistry, between functionalized polypeptide blocks and other polymer blocks. For instance, Deming and co-workers have synthesized different amphiphilic diblock copolypeptides containing a charged block and a hydrophobic block by using transition metal complexes as initiators.27 Kataoka and co-workers have prepared poly(ethylene glycol) (PEG)/polypeptide diblock copolymers by using amino-functionalized methoxy-poly(ethylene glycol)s (mPEG) as macro-initiators.12 Chen and co-workers have developed a series of polyester/polypeptide block copolymers by using amino-terminated polyester blocks as macro-initiators.29–34 Lecommandoux and co-workers have synthesized polybutandiene-polypeptite diblock copolymers by using monoamino-terminated polybutandienes as macro-initiators.35, 36 In addition, Bae and co-workers have developed a class of poly(L-histidine) (PLHis)-based block copolymers by coupling PLHis with other synthetic polymers.37

3. Stimuli-Sensitive Polypeptide-Based Materials

  1. Top of page
  2. Abstract
  3. 1. Introduction
  4. 2. Synthesis of Polypeptides
  5. 3. Stimuli-Sensitive Polypeptide-Based Materials
  6. 4. Stimuli-Sensitive Functionalized Polypeptides
  7. 5. Conclusions and Perspectives
  8. Acknowledgements
  9. Biographical Information

3.1. Micelles

Polymeric micelles are generally formed by self-assembly of amphiphilic polymers in aqueous solution.38 They have nano-sized amphiphilic structures composed of a hydrophobic inner core and a hydrophilic outer shell. The micellar cores are commonly consisted of hydrophobic segments or complexes of oppositely charged polyelectrolytes, which are shielded and stabilized by hydrophilic shells, such as an antifouling PEG shell. The core–shell nano-structures of polymeric micelles make them suitable for drug loading and delivery. The hydrophobic or ion complex cores of micelles facilitate the loading of hydrophobic drugs or ionic biopharmaceuticals, such as DNAs, RNAs, and proteins, while the hydrophilic antifouling shells improve the stability and biocompatibility of the micelles in circulation system in vivo. Additionally, nanometer- sized particles (10–200 nm) have been shown to passively accumulate in tumor sites due to the enhance permeation and retention (EPR) effect.5 Hence, nanomedicine approaches have emerged as the most popular methods for cancer therapy.39, 40 As a result, nanometer-sized polymeric micelles have received extensive investigation as drug delivery systems. Among them, micelles based on stimuli-sensitive synthetic polypeptides are interesting due to their stimuli-responsive self-assembly behaviors and conformation transitions as well as good biocompatibility and biodegradability.9–14 In recent years, polypeptide-based micelles comprising pH-sensitive polypeptide blocks, acid-labile polypeptide–drug conjugates, polypeptide containing ion complexes and dual stimuli-sensitive segments have been developed for controlled drug and gene delivery.

3.1.1. PEG-Poly(L-histidine) Block Copolymers

Drug nano-carriers may encounter different environmental stimuli, such as pH change and change in redox environment, after they are administrated in vivo. For example, weak acidic environments (pH 6.0–7.0) are formed in tumor sites, compared to the neutral pH (pH 7.4) of normal tissues and blood.5, 15 In addition, more acidic environments are formed as the nano-carriers are entrapped in endosomes (pH 5.0–6.5) and lysozomes (pH 4.0–5.0).5, 15 Therefore, pH-sensitive polymers have attracted considerable interest for their potential applications in anticancer and intracellular drug delivery.

In order to avoid enzymatic degradation within lysosomes, rapid endosomal/lysosomal drug release plays a key role in achieving ideal therapeutical efficiency. It has been pointed out that endosomal pH-triggered drug release may lead to a high intracellular drug dose and therefore result in improved efficiency in killing multi-drug resistance (MDR) tumor cells.41 Therefore, drug-loaded nanocarriers capable of responding to endosomal pH (5.0–6.5) may exhibit enhanced antitumor efficacy. Synthetic polypeptides containing ionizable pendant groups, such as poly(L-glutamic acid) (PLGlu), poly(aspartic acid) (PAsp), poly(L-lysine) (PLLys) and PLHis exhibit pH-induced hydrophilic-hydrophobic transitions at pH around their pKa values. Among the above polyelectrolytes, PLHis containing imidazole groups with a pKa of ∼6.0 has a sharp buffering capacity at endosomal pH.42 Additionally, hydrophilic PEG has been widely incorporated in biomaterials due to its good biocompatibility, non-immunogenicty and antifouling property, and PEG with relatively low molecular weights (MW < 50 000) can be excreted from the body through renal clearance.11, 12 Consequently, pH-sensitive micelles based on PEG–PLHis amphiphilic block copolymers have been intensively developed for drug delivery.

Bae and co-workers have developed a PEG-b-PLHis diblock copolymer by ROP of Nim-2,4-dinitrobenzene-L-histidine-NCA using n-hexylamine or isopropylamine as an initiator, followed by coupling monoamino-terminated PLHis with monocarboxylic PEG (Scheme2a).37 The pKa values of PLHis and PEG-b-PLHis were determined to be 6.5 and 7.0, respectively (Table1). The relatively higher pKa of PEG-b-PLHis was believed to be attributed to the hydration effect of hydrophilic PEG. The diblock copolymer formed core-shell micelles at pH above its pKa with the deprotonated PLHis segments as a hydrophobic core and PEG as a hydrophilic shell. The micelles showed an average diameter of ∼114 nm at pH 8.0. The PEG-b-PLHis micelles displayed a pH-induced destabilization behavior at pH below 7.4. With reducing the solution pH from 8.0 to below 7.2, the critical micelle concentration (CMC) of the copolymer increased obviously and the stability of the micelles decreased markedly, due to the protonation-deprotonation transition of the PLHis block. As doxorubicin (Dox) loaded micelles were cultured with A2780 cells, the in vitro uptake of Dox by A2780 cells was enhanced more than five times by reducing pH from 7.4 to 6.8, resulting in an increased cytotoxicity against A2780 cells at pH 6.8.43 The enhanced Dox cellular uptake and cytotoxicity at pH 6.8 was believed to be due to the acid-triggered release of Dox. After administration of the Dox-loaded micelles into mice, the Dox-loaded micelles showed markedly increased half life time in vivo compared to free Dox. After intravenous injection into nude mice bearing human ovarian A2780 tumor, the pH-sensitive drug loaded micelles exhibited significantly higher antitumor efficiency in comparison with free Dox.

Table 1. pKa values of some typical pH-sensitive peptide-based polymers.
Polymerp KaRef.
Poly(L-histidine) (PLHis)6.537
Poly(ethylene glycol)-b-poly(L-histidine) (PEG-b-PLHis)7.037
PEG-b-poly(L-histidine-co-L-phenylalanine) (PEG-b-P(LHis-co-LPhe))6.4–7.052
Poly(L-lactide)-b-PEG-b-PLHis (PLLA-b-PEG-b-PLHis)7.054
PEG-b-PLHis-b-PLLA-b-PEG6.2–6.863
PEG-b-poly(aspartic acid) (PEG-b-PAsp)4.8875
PEG-b-poly(L-lysine) (PEG-b-PLLys)9.5475
PAsp-b-PLLA6.7–7.256
PEG-b-poly(3-morpholinopropyl aspartamide)-b-PLLys (PEG-b -PMPA-b-PLLys)6.2 and 9.484
PEG-b-poly(N-(2-aminoethyl)-2-aminoethyl aspartamide) (PEG-b-(PAsp-DET))6.0 and 9.585
PEG-poly(2-aminopentyl-α,β-aspartamide) (PEG-P(Asp-AP))10.47157
Methylamine functionalized poly(L-glutamate)7.3255
Diethylamine functionalized poly(L-glutamate)6.5255
Diisopropylamine functionalized poly(L-glutamate)5.25255
Poly(L-glutamate)-g-oligo(2-aminoethyl methacrylate) (PLGlu-g -OAEMA)7.3240
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Scheme 2. Chemical structures of some representative polypeptide-based block copolymers that form stimuli-sensitive micelles.

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In order to improve the stability of the PEG-b-PLHis micelles, mixed micelles composed of PEG-b-PLHis and an amphiphilic PEG-b-poly(L-lactide) (PEG-b-PLLA) diblock copolymers have been subsequently developed.44 The mixed micelles maintained obvious pH-sensitivity provided that the weight ratio of PEG-b-PLLA within the blends is less than 50 wt%.45 With increasing the weight ratio of PEG-b-PLLA from 0 to 40 wt%, the destabilization pH of the mixed micelles was shifted from 7.2 to 6.6.44, 45 A formulation containing 25 wt% PEG-b-PLLA, i.e. mixed micelle (75/25), was selected to be the optimal formulation, for its transition pH (6.8) is close to the pH of the extracellular space of solid tumors. The mixed micelles (75/25) displayed a constant hydrodynamic diameter of around 129 nm in the pH range of 7.0–7.4. In contrast, the micelle size increased significantly to 195 nm as the pH was decreased from 7.0 to 6.8, indicating swelling of the mixed micelles caused by the protonation of the PLHis segments. In vitro cytotoxicity tests against human breast adenocarcinoma (MCF-7) cells indicated that the mixed micelles exhibited good biocompatibility.44 In vivo cytotoxicity test demonstrated that the micelles caused no apparent system toxicity in mice at concentrations up to 2400 mg kg−1.41 The Dox containing mixed micelles (75/25) showed a stable diameter of 183 nm at pH 7.4.46 In contrast, a significant increase in particle size was observed when the pH was reduced to 6.8, accompanying with a markedly accelerated drug release. A lyophilized micelles formulation has been prepared by using excipients.41 The reconstituted micelles showed a particle size, pH sensitivity and toxicity against MDR tumor cells comparable to the freshly made micelles. In order to impart active tumor-targeting ability to the pH-sensitive micelles, a tumor-targeting ligand, i.e., folic acid, was introduced into the surface of the micelles by mixing folate-conjugated PEG-b-PLHis (Fol-PEG-b-PLHis) with folate-conjugated PEG-b-PLLA (Fol-PEG-b-PLLA).47 It was found that the folate-conjugated pH-sensitive mixed micelles (Fol-PHSMMs) containing Dox exhibited significantly higher efficiency in killing MDR MCF-7 cells (MCF-7/DoxR) compared to free Dox and Dox-loaded micelles without folate. The enhanced efficacy in killing MDR tumor cells was proposed to be due to improved intracellular uptake via folate-receptor mediated endocytosis and subsequent endosomal escape. After injection into nude mice bearing Dox resistant MCF-7 (MCF-7/DoxR) xenografts through the tail vein, the Dox-loaded Fol-PHSMMs displayed markedly higher efficiency in suppressing tumor growth than either free Dox or Dox-loaded micelles without folate. The Dox accumulation level in tumors of the group treated with folate-conjugated micelles was twenty times higher than that of free Dox treated group and three times higher than that of the group treated with the micelles without folate. In addition, Dox-loaded Fol-PHSMMs have also been shown to have higher in vitro cytotoxicity against both wild type and MDR tumor cells than the folate-conjugated PEG-PLLA (Fol-PEG-b-PLLA) micelles.48 It was found that both micelles showed comparable cellular uptake levels of Dox due to the folate receptor-mediated encytosis, whereas cells treated with Fol-PHSMMs resulted in a widespread distribution of Dox in the cytosol, likely caused by the pH-triggered micelle destabilitzation and endosomal disruption. The rapid distribution of Dox within the tumor cells led to increased drug retention in both wild type and MDR tumor cells, and therefore resulted in an improved anti-MDR tumor efficacy.49, 50 The Dox-loaded Fol-PHSMM effectively suppressed the growth of MDR ovarian tumor in nude mice for 50 days by 3 intravenous injections at a dose of 10 mg kg−1 at a 3-day interval. In vivo optical imaging tests confirmed that the Cy5.5-labeled pH-sensitive micelles accumulated at tumor sites at 1 hour post-injection. A further in vivo study indicated that nude mice bearing metastatic 4T1 murine breast tumors treated with the Dox-loaded Fol-PHSMM exhibited marked suppression of tumor growth and no apparent metastasis at 28 days, compared to significant metastasis for the group treated with Dox solution in PBS, Dox-loaded PEG-b-PLLA micelles or Dox-loaded PEG-b-PLHis micelles without folate.51 In addition, the pKa of the PLHis-based copolymer was found to be influenced by introduction of hydrophobic unit, e.g., L-phenylalanine, to the PLHis segment.52 PEG-b-poly(L-histidine-co-L-phenylalanine) (PEG-b-P(LHis-co-LPhe)) diblock copolymers showed pKa values varying from 7.0 to 6.4, depending on the L-phenylalanine content within the P(LHis-co-LPhe) block.

3.1.2. Polypeptide–Polyester Block Copolymers

Biodegradable synthetic polymers, especially aliphatic polyesters and polycarbonates, have received extensive investigation in the past several decades, owing to their vast potential in biomedical applications, such as drug delivery and tissue engineering. In recent years, amphiphilic block copolymers containing a polypeptide block and a polyester block have attracted considerable attention,29–34, 53–58 due to their stimuli-responsiveness, good biocompatibility, adjustable biodegradability, tunable hydrophobicity, and the ability to incorporate functional groups.7 The above advantages make these materials interesting for biomedical applications.

As discussed in Section 3.1.1, micelles containing both polypeptide and polyester segments can be obtained by physically mixing two kinds of block copolymers, such as PEG-b-PLHis and PEG-b-PLLA.44 Additionally, micelles formed by polypeptide-polyester block copolymers have also been fabricated. A PLLA-b-PEG-b-PLHis triblock copolymer was fabricated via Michael-addition reaction between a thiol-terminated PLLA-b-PEG and a maleimide-functionalized PLHis.54 The triblock copolymer showed a pKa (∼7.0) close to that of PEG-b-PLHis. At pH above its pKa, flower-like micelles with a PLLA/PLHis mixed core and a PEG shell are formed, whereas swollen micelles with a PLLA core and a PEG/PLHis shell are obtained at lower pH. The triblock copolymer micelles showed a constant size of ∼80 nm at pH 7.2–7.4, but increased drastically with decreasing pH to 6.8. The release of Dox from the micelles was obviously accelerated with reducing pH from 7.4 to 6.0, attributed to the decrease in micellar stability caused by the ionization of PLHis segments. Some bioactive moieties, such as biotin and cell penetrating peptides (CPP),59 have been shown to markedly increase cellular uptake of drug nanocarriers. Nevertheless, nonspecificity of some bioactive molecules limits their applications in vivo. Due to its hydrophilic–hydrophobic transition at tumor extracellular pH (pH 6.0–7.0),5, 15 PLHis can be used as an “on–off” switch for selectively exposing these molecules at tumor sites. Biotin and CPP have been separately conjugated to the end of the PLHis block of PLLA-b-PEG-b-PLHis.60, 61 Mixed micelles comprising biotin conjugated PLLA-b-PEG-b-PLHis (PLLA-b-PEG-b-PLHis-Bio) and PEG-b-PLHis exhibited an acid-triggered avidin binding ability and improved cellular uptake by MCF-7 cells expressing biotin receptor at pH < 7.2.60 The pH-dependent avidin-binding behavior of the micelles is illustrated in Figure2. The biotin was tethered at the core-shell interface of the micelles at pH > 7.0 due to the deprotonation and hydrophobic nature of the PLHis spacer, leading to a low avidin-binding ability and lower uptake of the micelles by MCF-7 cells. In contrast, biotin was exposed to the surface of the micelles at 6.5 < pH < 7.0 caused by the partial protonation of PLHis, resulting in significant increase in cellular internalization of the micelles. At more acidic endosomal pH (<6.5), the drug loaded micelles destabilized, attributed to the further protonation of PLHis and the electrostatic repulsion between the PLHis segments. Mixed micelles based on a PLLA-b-PEG-b-PLHis linked with a CPP, i.e., transactivator of transcription (TAT) peptide, (PLLA-b-PEG-b-PLHis-TAT) were subsequently reported.61 The TAT-PLHis-conjugated micelles showed similar pH-dependent cell killing efficiency. The cytotoxicity was significantly increased as the pH decreased from 7.4 to 6.8, as compared with high cytotoxicities at all experimental pH for the micelles containing TAT linked through a PEG spacer. The pH-dependent cytotoxicity of the micelles with PLHis-linked TAT was believed to be due to the selective exposure of TAT at acidic pH values. Since slightly acidic microenvironments are formed in tumor sites,5, 15 it is envisioned that the PLHis spacer can be used to enhance tumor-targeting ability of the linked bioactive moieties. Additionally, PLHis/PLLA block copolymers with different architectures, such as PEG-b-PLHis-b-PLLA triblock copolymers and PEG-b-PLHis-b-PLLA-b-PEG tetrablock copolymers,62, 63 have been developed by combination of ROP of L-lactide (LLA) and coupling reactions. These block copolymer micelles showed similar pH-dependent micellar destabilization and accelerated drug release profiles at acidic pH.

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Figure 2. Schematic diagram depicting the central concept of pH-induced vitamin repositioning on the micelle. While above pH 7.0, biotin that is anchored on the micelle core via a pH-sensitive molecular chain actuator (polyHis) is shielded by PEG shell of the micelle; biotin is exposed on the micelle surface (6.5 < pH < 7.0) and can interact with cells, which facilitates biotin receptor-mediated endocytosis. When the pH is further lowered (pH < 6.5), the micelle destabilizes, resulting in enhanced drug release and disrupting cell membranes such as endosomal membrane. Reproduced with permission.60 Copyright 2005, American Chemical Society.

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In addition to coupling approaches, polypeptide–polyester block copolymers are largely synthesized via ROP of NCAs by using amino-functionalized polyesters as macroinitiators. Chen and co-workers have synthesized a series of block copolymers comprising polypeptide blocks and aliphatic polyester blocks,29–34 such as PLLA and poly(ϵ-carpolactone) (PCL), via ROP of NCAs by using mono- or diamino-functionalized polyesters as macro-initiators. Different bioactive moieties, such as RGD,30–32 biotin and sugar residues,64, 65 have been incorporated into the polypeptide segments through the pendent functional groups. The pH-dependent self-assembly behaviors of these block copolymers have been investigated.66 A series of PEG-b-poly(L-lactide)-b-poly(L-glutamic acid) (PEG-b-PLLA-b-PLGlu, Scheme 2b) were synthesized by ROP of γ-benzyl-L-glutamate NCA (BLG-NCA) using amino-terminated PEG-b-PLLA as a macro-initiator, followed by the deprotection of the benzyl groups.29 PLGlu shows a random coil-to-α-helix transition, associated with a hydrophilic-hydrophobic transition, as the pH decreases from neutral pH to the pH below the pKa (4.1) of the L-glutamic acid (LGlu) residues.4 The triblock copolymers formed micelles in water with the size and morphology depending on the pH and the PLGlu block length.66 For a triblock copolymer of PEG17-b-PLLA23-b-PLGlu60 (where the subscript numbers refer to the average number of the repeat unit for each segment), spherical micelles with an average diameter of 58 nm were observed at pH ≥ 4.5. The size of the spherical micelles decreased with decreasing pH to 3.9. Notably, when the pH was decreased to 3.2, rod-like micelles with a mean diameter of 28 nm were formed. pH-dependent 1H NMR tests revealed that the PLGlu segments were located in the outer shell part of the micelles at pH ≥ 4.5, whereas they were dehydrated and encapsulated by the PEG segments at pH 3.2. A three-layer structure was proposed for the rod-like micelles, which are composed of a PLLA core, a PLGlu intermediate layer and a PEG shell. The formation of the rod-like micelles was assumed to be attributed to the fact that the relatively short PEG block (degree of polymerization (DP) = 17) was not long enough to stabilize the spherical micelles at pH 3.2. On the other hand, only spherical micelles were detected within the experimental pH region for a triblock copolymer with a shorter PLGlu block (DP = 10). Additionally, the CMC of the triblock copolymer was found to increase with increasing the PLGlu block length or pH.

AB-type diblock copolymers with PLGlu or poly(aspartic acid) (PAsp) as the A block and PLLA as the B block have been synthesized by ROP of NCA using amino-terminated PLLA as a macroinitiator.55–58 Spherical micelles with a PLLA core and an ionizable polypeptide shell were formed by the diblock copolymers in water. The morphology and size of the micelles were found to be influenced by the relative block length and the pH.56 The pKa values (6.7–7.2) of PAsp-b-PLLA (Scheme 2c) were significantly higher than that of the pendant carboxylic groups of the aspartic acid (Asp) residues (3.9).4 The formation of hydrogen bonds between the PAsp segments was thoughted to be responsive for the increase in pKa.56 The size of the micelles increased markedly with increasing pH from 4 to 7 when the PAsp block is comparable or longer than the PLLA block. In contrast, when the PLLA block is dominant in the diblock copolymer, no obvious change in size was detected with increasing pH. In addition, in vitro cytotoxicity tests indicated that the polypeptide-polyester diblock copolymers displayed good biocompatibility.

3.1.3. Drug Conjugated Micelles via Acid-Labile Linkers

Amphiphilic PEG-polypeptide diblock copolymers, with anticancer drugs covalently conjugated to the polypeptide block, have advantages in anticancer drug delivery for their high stability in blood stream, increased solubility of hydrophobic drugs, less systematic toxicity, as well as improved accumulation at tumor sites owing to EPR effect.14, 15 Polypeptide–drug conjugates based on PLGlu and PEG-b-PAsp have been developed and entered clinical trails.14, 67, 68 It has been pointed out that the release of conjugated drugs from the polymer–drug conjugates after they are uptaken by tumor cells plays a crucial role in achieving desirable antitumor efficacy.16 Therefore, drugs have been linked to polymers via covalent bonds cleavable by physiologically relevant stimuli, such as acidic pH,15 reducing environment and enzymes.16 The stimuli-labile linkers may facilitate the release of the conjugated drugs from polymer backbones at desirable sites. Different polypeptide–drug conjugate micelles containing acid-labile linkers have been developed in recent years. Kataoka and coworkers have developed a PEG-poly(aspartate-hydrozone-adriamycin) (PEG-b-P(Asp-Hyd-ADR)) diblock copolymer, in which a anti-cancer drug, adriamycin (ADR), was conjugated to the PAsp block via a acid-cleavable hydrazone bond with a substitution ratio of 67.6 mol% (Scheme 2d).69 The block copolymer formed micelles in aqueous solution with the drug-conjugated PAsp segments as a core and PEG blocks as a shell. The micelles showed an average diameter of about 65 nm. The drug was tethered to the polypeptide backbone and loaded within the micelles at pH > 6.5, whereas the release of ADR was triggered at acidic pH. When cultured with human small cell lung cancer cell SBC-3, micelles trapped in endocytic compartments were observed by confocal laser scanning microscopy (CLSM), and cytotoxicity tests confirmed the antitumor efficacy of the ADR-conjugated micelles. In contrast to free ADR, the micelles displayed a time-dependent cytotoxicity, probably due to a delayed acid-triggered drug release process.70 Both in vitro tests using a multicellular tumor spheroid (MCTS) model and subsequent in vivo antitumor tests suggested the penetration of the ADR-conjugated micelles into tumor tissues.70 The ADR-conjugated micelles exhibited enhanced antitumor efficacy and lower toxicity compared to free ADR. A tumor-targeting polypeptide–drug conjugate was subsequently prepared by incorporation of folate to the PEG block of PEG-b-P(Asp-Hyd-ADR).71 Mix micelles composed of folate-functionalized PEG-b-P(Asp-Hyd-ADR) (Fol-PEG-b-P(Asp-Hyd-ADR)) and PEG-b-P(Asp-Hyd-ADR) were prepared and displayed an average size varying from 60 nm to 90 nm, depending on the substitution ratio of folate.72 Surface plasmon resonance (SPR) measurements confirmed the binding ability between the folate-conjugated micelles and folate-binding proteins (FBP). A folate conjugation ratio of 10% on the micelles was found to be enough to effectively bind FBP. The selective binding effect resulted in an enhanced intracellular uptake of the micelles and hence an increased efficiency in suppressing the growth of tumor cells.71 In vivo tests indicated that an increase in folate substitution ratio from 10% did not cause an increase in tumor accumulation of the micelles, attributed to the increased accumulation of micelles in liver. After intravenous injection into tumor-bearing mice, the folate-conjugated micelles with a folate conjugation ratio of 10% exhibited higher tumor suppression efficacy than either free ADR or micelles without folate. In addition, anticancer drugs have also been covalently conjugated to dendritic polypeptides via acid-labile linkers. Dox has been conjugated to a dendritic PLGlu containing biotin via hydrazone bonds by Gu and co-workers.73 The dendritic PLGlu-Dox conjugate formed spherical nanoparticles with an average diameter of around 50 nm in aqueous solution. The release of Dox from the nanoparticles was slow at pH 7.4, whereas it was dramatically accelerated at pH 5.0. In vitro cytotoxicity tests indicated that the Dox-conjugated nanoparticles showed an enhanced intracellular uptake and higher efficiency in killing of tumor cells.

3.1.4. pH-Sensitive Polyion Complex Micelles

Polyion complex (PIC) micelles are mainly formed by a charged amphiphilic block copolymer with an oppositely charged polymer or ionic biopharmaceuticals, such as DNA, RNA, and proteins. These materials are interesting for controlled delivery of ionic bioactive molecules, attributed to the fact that ionic biopharmaceuticals can be efficiently encapsulated in the micelles through electrostatic complexation.74 PIC micelles exhibit sensitivity to pH, salt and other competing polyelectrolytes. Since the commonly used biopharmaceuticals, such as DNA, RNA, and some proteins, are negatively charged at physiological condition, different positively charged polypeptides have been used for fabrication of PIC micelles. Kaotaka and co-workers have developed PIC micelles based on PEG-b-poly(L-lysine) (PEG-PLLys) diblock copolymers and DNA. Zeta (ζ) potential tests confirmed the formation of core-shell micelles with ionic complexes as a core and hydrophilic PEG segments as a shell.75, 76 The PIC micelles showed a constant diameter with the salt concentration varying from 0 to 300 mM. The shielding effect of the PEG shell led to increased stability of the complexes in the presence of nuclease or serum,76, 77 and reduced immunogenicity in vivo.78 The PIC micelles exhibited higher cellular uptake and gene transfection efficiency compared to a commercially available lipoplex system. After hydrodynamic injection into mice via limb vein, PIC micelles containing plasmid DNA (pDNA) displayed enhanced transgene expression than naked pDNA. After administration to tumor-bearing mice, PIC micelles containing pDNA encoding soluble vascular endothelial growth factor (VEGF) receptor-1 (sFlt-1) showed effective suppression of tumor growth. Active targeting ability has been imparted to the PIC micelles by introduction of cyclic RGD (cRGD),79 a moiety capable of recognizing αvβ3 and αvβ5 integrins overexpressed in tumor vasculature and malignant tumor cells.80 When cultured with HeLa cells expressing αvβ3 and αvβ5 integrins, the cRGD-conjugated PIC micelles showed higher transfection efficiency and enhanced cellular uptake than PIC micelles without cRGD. In contrast, for 293T cells without αvβ3 and αvβ5 integrins, both systems displayed comparable transfection efficiency. The selective increase in transfection efficiency suggested a receptor-mediated endocytosis mechanism. Additionally, PIC nano-carriers based on PLLys multiblock copolymers, graft copolymers, and hydrophobic segment containing triblock copolymers have also been developed.81–83

Due to the relatively high pKa (∼10) of PLLys,75 PLLys-based PIC micelles show no effective proton buffering capacity within the acidic endosomal pH range. Therefore, poly(3-morpholinopropyl aspartamide) (PMPA) with a pKa of 6.2 has been incorporated with PLLys to improve the buffering capacity of the material at endosomal pH.84 A PEG-b-PMPA-b-PLLys triblock copolymer was fabricated by sequential ROP of β-benzyl-L-aspartate NCA (BLA-NCA) and ϵ-(benzyloxylcarbonyl)-L-lysine NCA (ZLL-NCA), followed by aminolysis of benzyl ester of PBLA using 4-(3-aminopropyl)morpholine and deprotection of the benzyloxylcarbonyl groups of poly(ϵ-(benzyloxylcarbonyl)-L-lysine) (PZLL). PEG-b-PMPA-b-PLLys exhibited a two-step protonation process, and the triblock copolymer/pDNA complex formed a three-layer structure in aqueous solution, as illustrated in Figure3. At pH 7.4, the PLLys segments with a higher pKa (9.4) condensed pDNA in the inner core and the PMPA segments with a lower pKa (6.2) formed the intermediate layer. In comparison with PEG-b-PLLys/pDNA systems, the triblock copolymer-based PIC micelles displayed markedly increased transfection efficiency, due to the proton buffering capacity of PMPA segments at endosomal pH. Similarly, gene delivery systems showing two-step protonation process have been developed based on ethylenediamine-modified polyaspartamides,85 PLLys/PLHis graft copolymers,86 chitosan/PLLys graft copolymers,82 and aminoethyl-substituted PLHis.87 A N-(2-aminoethyl)-2-aminoethyl substituted PEG-polyaspartamide (PEG-P(Asp-DET), Scheme 2e) was synthesized by aminolysis of PEG-poly(β-benzyl-L-aspartate) (PEG-PBLA) using diethyenetriamine (DET).851H NMR test indicated the co-existence of α-aspartamide and β-aspartamide units within the poly(amino acid) backbone, caused by an intramolecular isomerization of aspartamide during the aminolysis. PEG-P(Asp-DET) showed a two-step protonation process with two pKa values of 6.0 and 9.5, respectively.85, 88 It was found that the pKa and buffering capacity of the N-substituted polyaspatamide were influenced markedly by the structure of the diamine side groups.85, 88 The two-step protonation process facilitates both DNA condensation at physiological pH (7.4) and endosomal escape at endosomal pH. DLS tests suggested that polyion complex (PIC) micelles consisting of PEG-P(Asp-DET) and pDNA revealed diameters of 70–90 nm with the N/P ratio ranging from 1 to 20. ζ potential measurements suggested that the ionic complex was shielded by a PEG shell. The PEG-P(Asp-DET)/pDNA micelles showed a pH-dependent membrane destabilization behavior, causing a relatively low cytotoxicity at pH 7.4 and an increased endosomal disruption at acidic environments.88, 89 In vitro transfection and cytotoxicity tests indicated that the PEG-P(Asp-DET)-based PIC micelles showed lower cytotoxicity and comparable or higher transfection efficiency to primary cells compared to either linear or branched polyethylenimine (PEI).85, 88, 89 The PIC micelles incorporated in calcium phosphate cement scaffolds were found to be released from the scaffolds in a sustained manner and successfully transfected surrounding cells.90 After administration of the scaffolds containing pDNAs expressing caALK6 and Runx2 to bone defects of mice, bone formation at the lower surface of the implants was observed for the mice treated with PIC micelles incorporated scaffolds, compared to no significant bone formation for the groups treated with commercially available linear PEI and FuGENE6. It was found that the low toxicity of P(Asp-DET) in vitro and in vivo is related to the rapid degradation of P(Asp-DET) segments caused by a self-catalytic process.91 P(Asp-DET) exhibited a decrease in cytotoxicity as the culture time increased, compared to no marked change in cytotoxicity of either linear PEI or DET-substitutetd poly(L-glutamide) (P(LGlu-DET)), which showed no obvious degradation in PBS during the experimental period. In order to further improve the stability of PIC micelles containing biopharmaceuticals, such as DNA and siRNA, PEG-P(Asp-DET) has been modified by different hydrophobic moieties, i.e., cholesterol and stearoyl groups.92, 93 The introduction of hydrophobic groups to the P(Asp-DET) block led to an increase in association force between the polymers and hence resulted in an improved micellar stability. Subsequent in vitro and in vivo studies suggested that the hydrophobically modified PIC micelles displayed increased transfection efficiency. Additionally, siRNA has also been conjugated to the poly(amino acid) backbone via disulfide bond, which is cleavable under intracellular reducing environment.94

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Figure 3. Chemical structure of PEG-b-PMPA-b-PLLys triblock copolymers and schematic illustration of the hypothesized three-layered polyplex micelles with spatially regulated structure. Reproduced with permission.84 Copyright 2005, American Chemical Society.

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The hydrophilic PEG shell of micelles plays an important role in improving biocompatibility and the blood circulation time of drug nanocarriers in vivo. On the other hand, the shielding effect of antifouling PEG segments may also lead to a decrease in cellular uptake and endosomal disruption.95, 96 Hence, drug and gene carriers with PEG segments cleavable under the microenvironments of tumor tissues and intracellular space have been fabricated.95–98 Detachable PEG blocks have been linked with P(Asp-DET) blocks via reduction-cleavable disulfide bonds or acid-cleavable hydrazone bonds.96, 97 It was found that the PIC micelles with a detachable PEG shell exhibited higher in vitro gene transfection efficiency than their counterparts with an undetachable PEG shell, likely due to an increased endosomal disruption and escape.96, 97 Stimuli-triggered desheilding systems have also been achieved by using ternary polyplex systems, in which ion complexes with excess positive charge are covered by a detachable negatively charged shell.99–101 A charge-reversal shielding polymer, cis-aconitic amide conjugated P(Asp-DET) (P(Asp-(DET-Aco))), has been synthesized by modification of P(Asp-DET) using an acid-labile cis-aconitic amide group.95, 100 After DNAs were condensed by excess cationic polymers, such as P(Asp-DET) and PLL, negatively charged P(Asp-(DET-Aco)) was incorporated to the surface of the ion complexes by elextrostatic interactions to form ternary polyion complexes. At acidic endosomal pH, the ourter layer detached from the ionic complexes due to the charge conversion of P(Asp-(DET-Aco)) caused by acidic hydrolysis of cis-aconitic amide. Besides their applications in gene delivery, pH-sensitive charge-conversional polypeptides have been investigated for controlled delivery of anticancer drugs and proteins.99, 102–104

Other examples of pH-sensitive shielding systems include carboxymethyl-substituted PLHis (CM-PLHis) and hydrophobically modified PLGlu,101, 105, 106 which both exhibited sharp response to endosomal pH. It is noteworthy that the ternary systems with a negatively charged surface exhibited no detectable cytotoxicity at physiological pH; on the other hand, excessive shielding polymers can also cause a decrease in transfection efficiency, due to a decrease in intracellular uptake.101, 106 Additionally, polyion complexes based on negatively charged polypeptides,107–110 including PEG-b-PAsp and PEG-b-PLGlu, and positively charged proteins or polysaccharides have been studied.

3.1.5 pH- and Temperature-Sensitive Micelles

Dual stimuli-sensitive polymers have attracted considerable attention due to their advantages in practical applications.3, 15 Especially, polymeric materials that sharply respond to physiologically relevant stimuli, especially pH and temperature, are interesting for controlled self-assembly and drug delivery.111 pH- and temperature-sensitive polypeptide-based copolymers are mainly fabricated by combination of pH-sensitive polypeptide segments with temperature-responsive synthetic polymers. In addition, dual stimuli-sensitive block copolymers may self-assemble into so-called ‘Schizophrenic’ micelles by varying temperature or pH.112 Very recently, pH- and temperature-sensitive polymers containing PLGlu, PLLys, and zwitterionic polypeptides have been synthesized and their micellization behaviors have been investigated.

Poly(N-isopropylacylamide) (PNIPAM) is one of the most widely studied temperature-sensitive polymers. It exhibits a sharp coil-to-globule transition at its low critical solution temperature (LCST) of around 32°C.18 Poly(N-isopropylacylamide)-b-PLGlu (PNIPAM-b-PLGlu, Scheme 2f) diblock copolymers have been synthesized via ROP of BLG-NCA by using monoanimo-terminated PNIPAM as a macroinitiator, followed by deprotection of the BLG groups.113–116 In order to obtain well-defined PNIPAM–polypeptide diblock copolymers, PNIPAM was synthesized via controlled radical polymerization,115–117 and ammonium chloride-functionalized PNIPAM was used as macroinitiators for ROP of NCAs.113, 118 PNIPAM-b-PLGlu showed a pH- and temperature-dependent micellization behavior in aqueous solution. Interesting “schizophrenic” micellization behaviors were observed by adjusting the pH and temperature, as illusrated in Figure4. The copolymer dissolves in PBS at neutral or alkaline pH and a temperature lower than its LCST. As the temperature is increased to above the LCST of PNIPAM, micelles with a PNIPAM core and an ionized PLGlu shell are formed, due to the dehydration of PNIPAM. In contrast, micelles with an α-helical PLGlu core and a PNIPAM shell are obtained with decreasing the pH to below the pKa (4.1) of LGlu residues.4 It was found that the LCST of the PNIPAM block showed a decrease as the PLGlu segment dehydrated at pH around the pKa of LGlu (4.1).113, 119 It is noteworthy that PLGlu does not exhibit a sharp hydrophilic–hydrophobic transition at endosomal pH, due to the relatively low pKa (4.1) of the pendant carboxylic groups and a gradual deprotonation process in a broad pH range.4, 120 To obtain polypeptides capable of responding to endosomal pH, Chen and co-workers synthesized a series of PNIPAM-poly[(L-glutamic acid)-co-(γ-benzyl L-glutamate)] (PNIPAM-b-P(LGlu-co-BLG)) diblock copolymers by partial deprotection of PNIPAM-b-PBLG.113 The existence of hydrophobic BLG within the polypeptide block led to an increase in the critical pH for the hydrophilic-hydrophobic transition of the polypeptide block. Interestingly, as the BLG content within the polypeptide block was higher than 30 mol%, the copolymers displayed sharp hydrophilic–hydrophobic transitions in response to a narrow pH change in the pH region of pH 5.5–7.4. Additionally, pH- and temperature-sensitive PLGlu-g-PNIPAM grafted copolymers have been synthesized by grafting monoamino-terminated PNIPAM to the pendant carboxylic groups of PLGlu.119 Amphiphilic aggregates were formed at neutral pH and a temperature above the LCST of PNIPAM. The size of the aggregates significantly decreased with reducing pH from 7.4 to 4.5, due to the gradual protonation and shrinkage of the PLGlu shell. In contrast, larger aggregates were detected with further decreasing pH to 4.2, caused by hydrophobic aggregation of PLGlu segments.

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Figure 4. Schematic illustration of the pH- and temperature-induced micellization behaviors of the PNIPAM-b-PLGlu diblock copolymer solution in PBS: micelles with an α-helical PlGlu core are formed at pH < 4 and 15 °C, whereas micelles with a PNIPAM core are formed at pH 7.4 and 37 °C.113, 114

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Cationic PNIPAM-b-PLLys diblock copolymers have also been synthesized via ROP of ZLL-NCA by using ammonium chloride-substituted PNIPAM as a macroinitiator by Chen and co-workers.118, 121 The chemical structure of PNIPAM-b-PLLys is shown in Scheme 2g. “Schizophrenic” micellization behaviors of the diblock copolymers were also observed by varying pH and temperature. PLLys exhibits a random coil-to-α-helix transition with increasing pH, and a hydrophilic–hydrophobic transition at pH around its pKa (∼10).75 Accordingly, PNIPAM-b-PLLys formed micelles with a hydrophobic PNIPAM core and a PLLys shell at lower pH and temperatures above its LCST, whereas micelles with a PLLys core and a PNIPAM shell are obtained at higher pH and lower temperatures. As shown in Figure5, environmental scanning electron microscopy (ESEM) measurements revealed that spherical PNIPAM-core micelles with a diameter of 150 nm were formed at pH 5.0 and 45 °C, while PLLys-core micelles with a diameter of 125 nm were observed at pH 12.5 and 25 °C. Similarly, zwitterionic PNIPAM- b-P(LLys-co-LGlu) diblock copolymers were synthesized by ROP of the mixture of ZLL-NCA and BLG-NCA using amino-terminated PNIPAM as a macroinitiator.122 The copolymers formed large aggregates in water at pH ranging from 6.0 to 10.0, due to strong intramolecular electrostatic interactions. In contrast, at pH < 6 or pH > 10, the electrostatic interactions were suppressed due to the deionization of LGlu or LLys residues, leading to a marked increase in solubility of the copolymer. In addition, a zwitterionic AB2 Y-shaped star polypeptide copolymer with PLLys as the A block and PLGlu as the B block was synthesized via “click” chemistry, and its pH-dependent “schizophrenic” micellization behavior was investigated.123 The zwitterionic copolymers may be interesting for some applications where dual stimuli-sensitivities at both acidic and alkaline conditions are needed.124

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Figure 5. ESEM image of spherical micelles formed by PNIPAM-b-PLLys (a) pH 5.0 and 45 °C, and (b) pH 12.5 and 25 °C, respectively. Reproduced with permission.118

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Other examples of pH- and temperature-sensitive polypeptide copolymers include copolymers based on PLGlu and poly(propylene oxide) (PPO) or poly(2-ethyl-2-oxazoline) (PEOZ).125–127 PLGlu-b-PPO-b-PLGlu triblock copolymers (Scheme 2(h)) were synthesized via ROP of BLG-NCA using bis(amino)-end-capped PPO as a macroinitiator, followed by deprotection of the BLG groups.125, 126 ‘Schizophrenic’ micellization behavior was observed for a PLGlu35-PPO33-PLGlu35 triblock copolymer with the PLGlu content of 78 wt%.125 At pH 10 and a temperature above the LCST of PPO (∼15 °C),128 PPO-core micelles with a diameter of ∼20 nm were formed, while PLGlu-core micelles were detected at pH 2 and 5 °C. Mixed micelles composed of PLGlu-b-PPO-b-PLGlu and a PEG-b-PPO diblock copolymer have been designed for controlled drug delivery.126 The hydrophobic PPO core was encapsulated by a PEG/PLGlu mixed shell at 25 °C and neutral pH. The release of loaded Dox from the micelles was found to be dependent on pH, temperature and the composition of the mixture. Drug release was accelerated at acidic pH (5.5). The increased release rate at acidic pH was proposed to be due to the formation of channels within the shell, caused by partial dehydration of PLGlu segments.

3.2. Vesicles

Polymer vesicles are nanometer-sized or micrometer-sized hollow spheres composed of a hydrophobic bilayer or interdigitated layer and hydrophilic internal and external shells.129–131 In contrast to the solid hydrophobic core of polymer micelles, polymer vesicles can envelop aqueous solution by their semipermeable membranes, making these materials ideal candidates for delivery of hydrophilic drugs and bioactive molecules.130 Additionally, hydrophobic drugs can also be loaded within the hydrophobic domain of the vesicle membrane. The size and permeability of the vesicle membrane can be tuned by adjusting the composition and chain length of the amphiphilic polymer. Especially, functionality and stimuli-sensitivity may be incorporated into the corona of vesicles, leading to intelligent systems for controlled drug delivery. Therefore, stimuli-sensitive vesicles based on different polypeptide block copolymers have been developed and tested for drug delivery applications.

3.2.1. Block Copolypeptides

A series of amphiphilic diblock copolypeptides with well-defined structures have been synthesized via successive ROP of NCAs by using transitional metal complexes as initiators by Deming and co-workers.27, 132 The diblock copolymers are constituted of a hydrophobic polypeptide block, such as poly(L-leucine) (PLLeu) and poly(L-valine) (PLVal), and a charged or hydrophilically modified polypeptide block, such as PLGlu, PLLys and poly(Nϵ-2-(2-(2-methoxyethoxy)ethoxy)acetyl-L-lysine) (P(EO2)LL). Hydrophobic PLLeu and PLVal are interesting for self-assembly due to their ordered conformation in aqueous solution.132 (P(EO2)LL)-b-PLLeu diblock copolymers with appropriate composition and chain length formed micrometer-sized vesicles (1–25 μm).133 It was found that the morphology of the assemblies was significantly influenced by the hydrophobic block content and the chain length of the whole copolymer. Spherical vesicles were formed at lower PLLeu contents and block lengths, while curved sheets were observed at higher PLLeu fractions and block lengths. As the PLLeu content was higher than 35 mol%, insoluble aggregates were formed. A pH-sensitive diblock copolypeptide, i.e., P((EO2)LL)160-b-P(LLeu0.3-co-LLys0.7)40, were fabricated by incorporation of LLys residues into the hydrophobic PLLeu block. As shown in Figure6, at a high pH (10.6), the diblock copolymer formed stable vesicles with a hydrophobic P(LLeu0.3-co-LLys0.7)40 layer and hydrophilic P((EO2)LL)160 inner and outer shells. Hydrophilic Fura-2 dye was enveloped within the vesicles during the self-assembly process. As the pH was reduced by addition of acid, the vesicle membranes were disrupted rapidly, resulting in the release of the encapsulated dye. It is noteworthy that the oligo(ethylene glycol) modified PLLys, i.e., P(EO2)LL, acts as a biodegradable PEG analogy and may exhibit good biocompatibility, making this material interesting for biomedical applications. It was found charged PLLys60-b-PLLeu20 (Scheme3a) and PLGlu60-b-PLLeu20 formed vesicles, and the size of the vesicles could be adjusted from nano-meter to micro-meter scale by a liposome-based extrusion technique.134 The vesicles were stable in 0.1 M PBS solution, and negatively charged PLGlu-b-PLLeu vesicles exhibited good stability in the presence of serum containing anionic proteins. In addition, vesicles have been developed based on an amphiphilic diblock copolypeptide consisting of PLLeu and polyarginine (PArg), which is a CPP analogy.135 The dextran-loaded vesicles showed enhanced cellular uptake of dextran with minimal cytotoxicity compared to free dextran.

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Figure 6. pH-responsive vesicles. a) Schematic drawing of P((EO2)LL)160-b-P(LLeu0.3-co-LL0.7)40 (denoted as K160p(L0.3/K0.7)40), its change in conformation with pH, and release of entrapped Fura-2 dye on pH change. b) Fluorescence emission as a function of excitation wavelength for Fura-2 dye (50 nM) entrapped in vesicles of K160p(L0.3/K0.7)40 in the presence of external calcium (5.0 mM) at pH 10.6 and 3.0. The frequency shift for maximum emission intensity at pH 3.0 is characteristic of calcium binding by Fura-2. Reproduced with permission.133 Copyright 2004, Nature Publishing Group.

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Scheme 3. Chemical structures of some typical polypeptide-based block copolymers that form stimuli-sensitive vesicles.

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Vesicles based on block copolypeptides containing other hydrophobic polypeptide blocks have also been reported. Jing and co-workers have synthesized a series of PLLys-b-poly(L-phenylalanine) (PLLys-b-PLPhe, Scheme 3b) by successive ROP of ZLL-NCA and L-phenylalanine NCA (LPhe-NCA) using n-hexylamine as an initiator, followed by deprotection of the ϵ-benzyloxycarbonyl groups.136 Interestingly, the diblock copolypeptides with relatively shorter PLPhe blocks could be directly dissolved in water and spontaneously form giant vesicles. Atom force microscopy (AFM) measurements confirmed the formation of vesicles, as shown in Figure7A. The height of the central part of the vesicle was lower than that of the peripheral part, due to the collapse of the hollow vesicles. Confocal laser scanning microscopy (CLSM) tests indicated that a hydrophilic dye, Rhodamine B, could be encapsulated within the interior aqueous compartments of the vesicles. 1H NMR measurements in D2O suggested that the PLPhe segments were surrounded by the hydrophilic PLLys blocks. Similar to the aforementioned P(EO2)LL-b-PLLeu system, a higher content of the rigid PLPhe, e.g., >35 mol%, led to a water-insoluble diblock copolymer. The vesicle size also exhibited a dependence on the initial polymer concentration and the solution pH. The average vesicle size decreased as the pH was increased from 2.5 to 9, due to a gradual decrease in electrostatic repulsion between the PLLys segments. Protein-encapsulated PLLys-b-PLPhe vesicles were prepared by direct dissolving the copolymer in PBS solution containing FTIC-BSA or carbonylated hemoglobin (CO-Hb) (Figure 7B).137 The encapsulation efficiency was found to be significantly affected by the interactions between the positively charged PLLys and the proteins. At a pH (5.8) slightly above the isoelectric point (PI) of FITC-BSA (PI = 4.8), negatively charged BSA was absorbed on the surface of the positively charged vesicles. In contrast, at a pH (3.8) below the PI, BSA was encapsulated within the inner compartments of the vesicles, due to the electrostatic repulsion between PLLys segments and positively charged BSA. Similarly, CO-Hb was enveloped in the vesicles at a pH (5.8) slightly below its PI (6.68). The encapsulation efficiency of CO-Hb was about 40 wt% and the loading content was 32 wt%. It was found that CO-Hb encapsulated within the vesicles retained its bioactivity and showed enhanced stability as compared with free CO-Hb. Aqueous solution containing O2 diffused into the interior compartments, causing the conversion of CO-Hb to oxygen-binding hemoglobin under irradiation of visible light.

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Figure 7. A) AFM image of vesicles formed in a PLLys-b-PLPhe aqueous solution at pH 3.5: A-a) AFM height image, A-b) AFM phase image, A-c) height profile along the line indicated in (A-a). B) ESEM image of CO-Hb-encapsulated PLLys-b-PLPhe vesicles. A,B) Reproduced with permission.136, 137 Copyright 2007 and 2009, respectively, American Chemical Society.

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An anionic PLGlu-b-PLPhe diblock copolymer has also been synthesized through consecutive ROP of BLG-NCA and LPhe-NCA using n-butylamine•HCl as an initiator.138 At pH > 6.0, vesicles with an average diameter of about 150 nm were formed by dissolving the diblock copolymer in deionized water. With decreasing pH from 6.0 to 5.0, the size of particles increased due to the gradual aggregation of the particles, and larger aggregates were formed at pH below 5.0. The self-assembly behavior of diblock copolypeptides containing a hydrophobic polypeptide block without an ordered secondary conformation, e.g., polyglycine (PGly), has also been investigated. The vesicle membrane formed by PLLys-b-PGly diblock copolymers was found to exhibit a high fluidity and flexibility.139

In addition to the diblock copolypeptides, amphiphilic triblock copolymers comprising a rigid hydrophobic central block flanked by two charged blocks, i.e., PLLys-b-PBLG-b-PLLys, have been developed. The PLLys-b-PBLG-b-PLLys triblock copolymers were prepared via sequential ROP of BLG-NCA and ϵ-Boc-L-lysine NCA (BocLL-NCA) by using 1,6-diaminohexane as a difunctional initiator under high vacuum, followed by selective deprotection of the Boc groups.140 Interestingly, in comparison with a relatively small composition range for the formation of vesicular structures by diblock copolypeptides containing a rigid hydrophobic polypeptide block (e.g., <35 mol%),133, 136 vesicles were formed by the triblock copolypeptides with the rigid hydrophobic block content of 19–74 mol%. The difference was believed to be attributed to the fact that a monolayer membrane can be easily formed by the triblock architecture, whereas an antiparallel orientation is needed for the formation of the vesicle bilayer membranes by amphiphilic diblock copolymers.140 It was found that an increase in the content of the middle PBLG block led to a more robust vesicular structure. The vesicles displayed a pH- and temperature-responsive structure conformation. The vesicle size decreased with reducing pH from 7.4 to 11.7 at 25 °C, due to the conformation change of the PLLys block from extended random coil to compact α-helix. In addition, the size showed an increase with increasing temperature from 25 to 37 °C at pH 11.7, resulted from the conformation transition of the PLLys block from α-helix to an elongated β-sheet conformation. When mixed with pDNA, the triblock copolypeptides also formed vesicles with higher size compared to the vesicles without pDNA, and it was found that pDNA was both complexed on the PLLys corona and encapsulated within the vesicles.

Vesicles based on a zwitterionic diblock copolypeptide have also reported. A zwitterionic PLGlu-b-PLLys (Scheme 3c) diblock copolymer with equal lengths of PLGlu and PLLys was synthesized via successive ROP of trifluoroacetyl-L-lysine NCA (TFA-LL-NCA) and BLG-NCA, followed by deprotection of TFA and BLG groups by using KOH.141 Spherical vesicles with a uniform diameter of 110 nm were formed at acidic pH (pH < 4) with α-helical PLGlu blocks as the intermediate hydrophobic layer and PLLys as the inner and out shells, whereas vesicles with a size of 175 nm composed of reverse inside-out components were formed at alkaline pH (pH > 10). At near neutral pH, the diblock copolymers formed precipitates due to the electrostatic interactions between the oppositely charged segments. The formation of vesicles was believed to be related to the presence of a rod-like conformation in the hydrophobic segment, leading to a low interfacial curvature and a hollow structure.

3.2.2. Polypeptide–Polyester Block Copolymers

Some polypeptide-polyester block copolymers containing an ionizable polypeptide block and a hydrophobic polyester have been shown to form pH-sensitive micelles in water, as discussed in Section 3.1.2.54–58, 66 In addition, vesicular structures have also been observed for some polypeptide-polyester diblock copolymers with short polypeptide blocks. A poly(lactic-co-glycolic acid)-b-PLGlu (PLGA-b-PLGlu) diblock copolymer with a short PLGlu block (DP = 15, f = 18 wt%) was prepared by coupling carboxylic group functionalized PLGA with amino-terminated PBLG, followed by deprotection of the benzyl groups (Scheme 3d).142 The diblock copolymer exhibited a pH-dependent self-assembly behavior. At a pH (3.0) below the pKa (4.1) of LGlu, the diblock copolymer formed large hydrophobic aggregates, due to the hydrophobic nature of both blocks. As the pH was increased to 5.0, spherical micelles with a mean hydrodynamic diameter of 81.5 nm were observed. The micelles were believed to be composed of a hydrophobic PLGA core and a relatively hydrophilic PLGlu shell. With further increasing the pH to 7.0, uniform vesicular structures with an average size of about 160 nm were obtained. The vesicles were constituted of a PLGA intermediate layer and ionic PLGlu inner and outer shells. The formation of vesicles was believed to be driven by an increase in the degree of ionization of PLGlu and the electrostatic repulsion between the short polypeptide segments. As the pH was further increased to 9.0, the vesicle size increased, likely due to the further increase in the electrostatic repulsion between PLGlu segments caused by the further deprotonation of PLGlu in water.120

3.2.3. Polypeptide-Polycarbonate Block Copolymers

Besides polyesters, polycarbonates are another important class of biodegradable synthetic polymers. Amphiphilic block copolymers composed of a polypeptide block and a polycarbonate block have been reported recently. Poly(trimethylene carbonate)-b-PLGlu (PTMC-b-PLGlu, Scheme 3e) diblock copolymers with PLGlu contents of 39 wt% and 46 wt%, respectively, were synthesized by ROP of BLG-NCA using a monoamino-functionalized PTMC, followed by deprotection of the BLG groups.143, 144 Vesicles comprising a PTMC intermediate layer as well as inner and outer PLGlu shells were prepared by either direct dissolution or solvent displacement method (nano-precipitation). The vesicles exhibited an almost constant size at pH not less than 7, whereas the size decreased with reducing pH from 7, due to a decrease in the electrostatic repulsion between the PLGlu blocks and a random coil-to-α-helix transition of the PLGlu chains. Additionally, the vesicles exhibited high stability against nonionic surfactant and nonpermeability to water, due to the high thickness of the vesicle membrane (30 nm). In contrast, the vesicles were rapidly disrupted in the presence of enzyme. The encapsulation of an ionizable anticancer drug, i.e., Dox, by the PTMC-b-PLGlu vesicles was found to be significantly influenced by the loading pH, likely due to the ionization of Dox at pH below its pKa (8.25).145 When loaded at pH 7.4, positively charged Dox was partially absorbed in the PLGlu shell; in contrast, when loaded at pH 10.5, neutral Dox was encapsulated inside the vesicles. The Dox-loaded vesicles displayed a loading content and loading efficiency of 47 wt% and 67 wt%, respectively, at pH 10.5. It was found that the release of Dox from the vesicles was affected by both pH and temperature. An obviously faster release profile was observed at pH 5.5 in comparison with the release pattern at pH 7.4. The higher release rate at pH 5.5 was thought to be resulted from a higher hydrophilicity of Dox at lower pH. Additionally, an increase in temperature from 25 to 45°C resulted in an increase in drug release rate, probably due to an increase in permeability and mobility of the membrane.

3.2.4. Polydiene-Polypeptide Block Copolymers

Polypeptide-based block copolymers containing a polyvinyl block or a polydiene block have also been fabricated, and their stimuli-sensitive self-assembly behaviors have been investigated. Gallot et al. and Klok et al. have investigated the self-assembly of different polystyrene–polypeptide block copolymers with a rod–coil structure.146–148 Especially, polypeptide-based block copolymers containing a polydiene block have attracted considerable interest, attributed to the fact that the self-assembled nanoaggregates can be further developed into shape-persistent nanostructures by photocrosslinking of the 1,2-vinyl bonds remained within the polydiene blocks.35 In two separate studies, polybutadiene-b-PLGlu (PB-b-PLGlu, Scheme 3f) diblock copolymers were synthesized by anionic polymerization of butadiene and subsequent ROP of BLG-NCA.35, 149 It was found that the diblock copolymers formed spherical micelles or vesicles after direct dissolution in water, depending on the composition and overall chain length.149, 150 Generally, spherical micelles should form when the fraction of the hydrophilic block is much higher than that of the hydrophobic block.151 The PB-b-PLGlu micelles and vesicles exhibited almost constant sizes at pH between 7 and 12, whereas their sizes showed a marked decrease as the pH was decreased from 7, due to the conformation transition of PLGlu.36 The vesicles formed by PB40-b-PLGlu100 showed a hydrodynamic radii ranging from 100–150 nm, depending on the pH.152 After irradiated by UV for 1h, the vesicles showed an enhanced stability and a reduced swelling ratio in THF, due to the covalent crosslinking of the vesicle membrane.

In addition to PB-b-PLGlu, positively charged PLLys has also been incorporated with diene polymers to fabricate amphiphilic diblock copolymers. Savin et al. and Schlaad et al. have prepared PB-b-PLLys diblock copolymers and investigated the solution properties of the vesicles formed by PB-b-PLLys.153, 154 It was found that the vesicles responded to both pH and temperature. The pH-dependent size change of the aggregates exhibited a manner opposite to the PB-b-PLGlu systems. At higher pH, the vesicles showed a smaller size and chains were packed more compactly at the core–corona interface.154, 155 With decreasing pH to less than 5.5, the size increased significantly. On the other hand, at a pH (10.9) slightly above the pKa (10.5) of the LLys residues,4 the vesicle size increased obviously as the temperature was increased from 25 °C to above 40 °C, caused by an α-helix-to-β-sheet transition.152 Lecommandoux and co-workers have synthesized a series of polyisoprene-b-PLLys (PI-b-PLLys) diblock copolymers via ROP of ZLL-NCA using monoamino-functionalized PI, followed by deprotection of the ZLL groups.155 It was found that the formation of micelles and vesicles by the diblock copolymers was dependent on the molar ratio of the polypeptide block to the PI block. A general conclusion was proposed that, to form a vesicular structure, the molar ratio of the charged polypeptide block within the PI-b-PLLys diblock copolymers is usually lower than 65 mol%.155

3.2.5. Polyion Complex Vesicles

A series of polyionic complex (PIC) micelles based on positively charged PEG-polypeptide block copolymers and negatively charged biopharmaceuticals, such as DNA, RNA and some proteins, have been designed for gene and protein delivery, as discussed in Section 3.1.4. In addition to the PIC micelles containing biopharmaceuticals, PIC vesicles based on two oppositely charged polypeptide block copolymers have also developed. These vesicular structures are composed of a PIC intermediate layer and hydrophilic inner and outer shells. Kataoka and co-workers have prepared a PIC vesicle with a diameter up to 10 μm by negatively charged PEG-PAsp and positively charged PEG-poly(2-aminoethyl-α,β-aspartamide) (PEG-P(Asp-AE)) or PEG-poly(5-aminopentyl-α,β-aspartamide) (PEG-P(Asp-AP)).156 It was found that the formation of PIC vesicles was influenced by both the charge ratio and relative block length. PEG-P(Asp-AP) and PEG-PAsp have pKa values of 10.47 and 4.88, respectively, and hence are almost equally charged at physiological pH.157 The morphology of the PIC vesicles formed by PEG-P(Asp-AP) and PEG-PAsp displayed a pH-dependence. Stable vesicles were formed at pH > 5.7; in contrast, the vesicles were disassociated into smaller particles as the pH was decreased to 5.7 or less, caused by a decrease in association force of the complexes. Dextran encapsulated vesicles were prepared by simple mixing of PEG-P(Asp-AP) and PEG-PAsp in the presence of FITC-dextran (FITC-Dex). The FITC-Dex encapsulated vesicles showed good stability even in the presence of 10% fetal bovine serum. The permeability of the PIC vesicles was found to be dependent on both pH and the MW of the guest macromolecules. The permeation of dextran labeled with tetramethylrhodamine isothiocyanate (TRITC-Dex, Mn = 70 000) from the vesicles was significantly faster at pH 5.8 than at pH 6.2–7.4, resulted from a reduced association force of the PIC layer at acidic pH. Additionally, a smaller guest molecule led to a higher permeation rate. Myoglobin (Mb) encapsulated vesicles with sizes ranging from 500 nm to 5 μm were prepared by mixing of an PEG-PAsp aqueous solution containing Mb with an PEG-P(Asp-AP) aqueous solution.158 Mb encapsulated inside the vesicles retained its bioactivity, and reversible oxygenation/deoxygenation of Mb was achieved in the presence of trypsin in the outer medium. In addition, polypeptide-based hollow capsules have also been prepared by layer-by-layer (LbL) assembly of PLLys and PLGlu on colloidal supports, followed by crosslinking of the film via ‘click’ chemistry.159

3.2.6. Polypeptide-Based A2B Lipid Mimetics

A2B amphiphilic copolymers comprising two lipophilic A segments and a hydrophilic B block may facilitate the formation of vesicular structures, because the hydrophobic volume is maximized and hence the formation of a bilayer structure is favored.160 Very recently, pH-sensitive A2B lipid mimetics with PLGlu as the B block and different lipophilic moieties, including octadecane, cholesterol and polyhedral oligomeric silsesquioxane, as the A segment have been developed via thiol-alkyne “click” chemistry.160, 161 All the above lipid mimetics were found to form vesicular structures. CD, DLS and SLS tests indicated that the vesicle size increased as the pH increased from 4 to 10, while the aggregation number of the vesicles showed an almost constant value within the experimental pH range, suggesting that the pH-induced size change of the vesicles was mainly caused by an α-helix-coil transition of PLGlu and a change in chain packing. Additionally, A2B lipid mimetics with hydrophobic polypeptide or polyester as the A blocks have been developed.162, 163

3.2.7. pH- and Temperature-Sensitive Vesicles

Vesicles formed by pH- and temperature-sensitive copolymers are interesting for controlled self-assembly and drug delivery, because the formation and disruption of the vesicular structures can be controlled by varying pH or temperature. In addition, “schizophrenic” vesicles may be formed by some pH- and temperature-responsive block polymers with appropriate polymer architectures and block length. Recently, vesicles based on pH- and temperature-sensitive polypeptide copolymers have been investigated. Savin and co-workers prepared a series of PPO-b-PLLys diblock copolymers via ‘click’ reaction between azide-terminated PPO and acetylene-terminated PLLys.164 The diblock copolymers displayed a random coil-to-α-helix transition at pH (7–8) lower than the pKa of the PLLys homopolymers (∼10),75 owing to segmental charge repulsion. At pH 3 and a temperature (25 °C) higher than the LCST (∼15 °C) of PPO, vesicles with diameters ranging from 48 to 100 nm were formed by PPO-b-PLLys with the PLLys weight contents of 34 wt%–76 wt%, while spherical micelles were observed for the sample with the PLLys content of 92 wt%. In contrast, as the pH was increased from 3 to 7 at 0 °C, aggregates with an α-helical PLLys core and a PPO corona were detected. It was found that the presence of a triazole ring did not show significant effect on the morphology of the aggregates, in contrast to an increase in interfacial curvature for the systems formed by polypeptide-driven self-assembly.165 The loading of a water-soluble anticancer drug, Dox•HCl, by the PPO-b-PLLys micelles and vesicles were compared. At pH 6.0, Dox•HCl encapsulated in vesicles exhibited a slow and sustained release profile with an initial burst, in comparison with a rapid release pattern of Dox•HCl from the PPO-b-PLLys micelles. This fast release of Dox•HCl from the PPO-b-PLLys micelles was assumed to be attributed to the fact that the drug was mainly loaded within the hydrophilic shell. In addition to the above positively charged systems, anionic pH- and temperature-sensitive vesicles have also been investigated. Lin and co-workers have found that the formation of vesicular or micellar structures by a PLGlu-b-PPO-b-PLGlu triblock copolymer could be controlled not only by polymer composition but also by pH.166 Only spherical micelles were observed for PLGlu-b-PPO-b-PLGlu with the PLGlu content of 74.5 mol%, whereas only vesicles were formed for a sample with the PLGlu content of 34 mol% at both acidic (pH 4.1) and basic pH (pH 11.5). On the other hand, for a sample with an intermediate PLGlu content, spherical micelles were observed at pH 11.5, whereas a mixture of vesicles and micelles were detected at pH 4.1. The self-consistent field theory (SCFT) simulation results suggested that the morphology change from micelles to vesicles was driven by a decrease in the hydrophilic-hydrophobic interfacial area caused by a decrease in the hydrophilic PLGlu block length or by the pH-induced coil-to-helix transition of the PLGlu block.

Tertiary amine-modified polymethacrylates, such as poly(2-dimethylamino)ethyl methacrylate) (PDMAEMA) and poly(2-diethylamino)ethyl methacrylate) (PDEAEMA), are unique pH- and temperature-sensitive polymers that exhibit LCST behaviors at pH slightly above the pKa of the pendant tertiary amines.167 For instance, PDMAEMA has a pKa of 7.0–7.8, and deionized PDMAEMA exhibits LCST of 32–46 °C, depending on the MW.3, 167 pH- and temperature-responsive vesicles based on PDMAEMA-b-PLGlu diblock copolymers have been reported recently.168, 169 The PDMAEMA-b-PLGlu (Scheme 3g) diblock copolymers were synthesized by coupling PDMAEMA and PBLG via Cu(I)-catalyzed azide-alkyne “click” chemistry, followed by deprotection of the benzyl groups.168, 169 The diblock copolymers exhibited PI values ranging from 4 to 8.5. At pH 6 and 25 °C, vesicles with a size of 300 nm were formed by a diblock copolymer, i.e., PDMAEMA85-b-PLGlu77, due to the electrostatic interactions between PDMAEMA blocks and PLGlu blocks. The vesicles were stabilized by excessive amino groups. As the pH decreased from 6 to 4, the vesicle size decreased to 138 nm. In addition, temperature-induced formation of vesicles by the diblock copolymers was also observed at a pH (pH 11) higher than the pKa of PDMAEMA. It was found the formation of micelles and vesicles driven by the dehydration of the PDMAEMA block was dependent on the relative block length. At pH 11 and 60 °C, vesicular stuctures were formed by the diblock copolymers with PLGlu contents of 26 wt% and 43 wt%, whereas micelles were obtained for the diblock copolymer with a PLGlu content of 64 wt%.

3.3. Crosslinked Nanoparticles

Stimuli-sensitive crosslinked nanoparticles, such as crosslinked micelles and hydrogel nanoparticles, have unique advantages including high stability, high drug loading capacity, and ability to respond to environmental stimuli.170–172 Hydrogel nanoparticles, i.e., nanogels, are swollen nanometer-sized networks comprising hydrophilic or amphiphilic polymers.170, 171 Nanogels may be composed entirely of a polymeric network or a core-shell structure with a hydrogel core or shell.171 Stimuli-sensitive nanogels with an hydrophilic and biocompatible shell, such as stimuli-dependent swellable crosslinked micelles, have received much attention due to their smart swelling–deswelling transitions and improved biocompatibility. Drugs can be loaded in the temporarily stable nanocarriers and the release of the drugs can be triggered by environmental stimuli at desirable sites. In the past decade, polypeptide-based crosslinked nanoparticles, such as crosslinked micelles and nanogels, which respond to external stimuli, such as pH, reducing environment, and dual stimuli, have been developed for different drug delivery applications. These systems have been fabricated through various methods including crosslinking of preformed polymer chains or self-assembled nanoaggregates, precipitation polymerization, and one-step ring-opening polymerization.

3.3.1. Reduction-Sensitive Crosslinked Nanoparticles

Nanoparticles crosslinked by covalent bonds capable of responding to intracellular microenvironment, in particular the acidic endosomal pH and the reducing environment of the cytoplasm, are interesting for intracellular drug delivery.5, 17 The development of reduction-sensitive nanocarriers is commonly achieved by incorporation of disulfide bonds to the drug delivery systems, due to selective cleavage of the disulfide bonds in the cytoplasm. The disulfide bonds are stable in the relative oxidation environment in vivo, such as body fluids and extracellular space, resulted from a low concentration (2–20 μM) of glutathione (GSH).17 In contrast, the disulfide bonds can be cleaved in the relative reducing environment of the cytoplasm due to a relatively high GSH concentration (0.5–10 mM). Consequently, disulfide-crosslinked nanoparticles based on polypeptides have also been developed for intracellular drug and gene delivery.173–175

Reversible shell crosslinked micelles based on a series of poly(L-cysteine)-b-PLLA diblock copolymers (PLCys-b-PLLA, Scheme4a) have been synthesized.173 PLCys-b-PLLA was synthesized via ROP of β-benzyloxycarbonyl-L-cysteine N-carboxyanhydride (ZLC-NCA) by using amino-terminated PLLA as a macro-initiator, followed by deprotection of the benzyloxycarbonyl groups using HBr. Micelles with a PLLA core and a PLCys shell were firstly formed in the presence of DTT, and then shell crosslinked micelles containing disulfide bonds were obtained by removing DTT and aerial oxidation. Ellman's assay suggested that less than 6% free thiols remained after oxidation. The decrosslinking of the micelles was achieved by addition of DTT. DLS and ESEM measurements revealed that the particle size increased slightly from 41.7 nm to 55.1 nm after adding DTT, and decreased to 47.1 nm again following removal of DTT, suggesting a reversible crosslinking–decrosslinking process of the micelles. Additionally, no intermicellar crosslinking was observed by DLS and ESEM tests. A hydrophobic model drug, rifampicin, was loaded into the shell crosslinked micelles by self-assembly of PLCys-b-PLLA in the presence of rifampicin, followed by aerial oxidation.176 The drug-loading content and loading efficiency were 15.0% and 17.5, respectively. The drug-loaded crosslinked micelles showed a size (∼65 nm) slightly higher than the particles without drug. A faster release profile was observed in the presence of DTT, due to decrosslinking of the particles caused by cleavage of the disulfide bonds by DTT. In addition, a model protein containing a free cysteine residue, i.e.; BSA, was successfully conjugated to the diblock copolymer via oxidation, and the release of BSA from the copolymer was triggered by addition of reducing agent.

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Scheme 4. Chemical structures of some representative polypeptide-based block copolymers and crosslinkers that are used for the preparation of crosslinked micelles and nanogels.

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Reversible-core crosslinked micelles containing disulfide bonds have also been developed. Kataoka and co-workers have prepared a series of core-crosslinked polyion complex (PIC) micelles via disulfide bonds through mixing of thiolated PEG-b-PLLys and PAsp, followed by crosslinking of the micellar core via aerial oxidation.174 The thiolation of PEG-PLL was performed via the coupling reaction between PEG-b-PLLys and N-succinimidyl 3-(2-pyridyldithio)propionate (SPDP) or 2-iminothiolane (IM).177 The IM-conjugated PLLys block has a higher charge density than the SPDP-modified PLLys block, due to the introduction of positively charged imino moieties. The chemical structure of a thiolated PEG-b-PLLys (denoted as PEG-b-P(LLys-MP)) obtained by treating the SPDP-modified PEG-b-PLLys using DTT is shown in Scheme 4b. The core-crosslinked PIC micelles showed an almost constant diameter with the salt concentration ranging from 0 to 0.5 M, in contrast to a significant decrease in stability of the parent uncrosslinked PIC micelles.174 On the other hand, the disulfide crosslinked micelles were found to be dissociated rapidly in the presence of a reducing agent, DTT, indicating a reduction-induced cleavage of the disulfide bonds. It is noteworthy that the reductive stability of disulfide bonds can be affected by the substitution groups close to the disulfide bonds.178 In subsequent studies, disulfide core-crosslinked PIC micelles based on thiolated PEG-PLLys and negatively charged biopharmaceuticals, including antisense oligonucleotide (ODN), pDNA, and siRNA, were developed.177, 179, 180 It was found that the core-crosslinked PIC micelles with thiolation degrees of 10 and 26% displayed sizes comparable to that of the uncrosslinked micelles.179 The core-crosslinked PIC micelles exhibited a high stability in the presence of a competing polyanion, i.e., poly(vinyl sulfate) (PVS), and showed a markedly enhanced resistance to nuclease compared to free ODN and uncrosslinked micelles. The release of loaded biopharmaceuticals from the crosslinked PIC micelles was triggered by addition of reducing agents, such as GSH. A higher GSH concentration and a lower crosslinking density resulted in a faster release rate. The disulfide crosslinked PIC micelles containing pDNA with a size of ∼100 nm showed enhanced in vitro transfection efficiency than uncrosslinked micelles.177 This was believed to be due to a higher stability of the crosslinked micelles in culture medium, leading to an increased cellular uptake.181 In addition, the gene transfection efficiency was also influenced by a balance between the charge density and thiolation ratio. A freeze-dried formulation of the core-crosslinked PIC micelles containing pDNA was subsequently developed.182 The micelles with a thiolation degree higher than 13% retained the original size and shape and showed comparable gene transfection efficiency after a freeze-drying/reconstitution process. After intravenous injection of the crosslinked micelles with a thiolation degree of 37% into mice via the orbital vein, a gene expression was observed in the liver, compared to no gene expression detected for the group treated with naked pDNA. Crosslinked PIC micelles conjugated with a targeting ligand, i.e., cRGDfk, have also been prepared by thiolated cRGDfk-PEG-PLLys and pDNA.181 The introduction of cRGDfk ligand led to an increase in transfection efficiency against HeLa cells expressing αvβ3 integrin receptors. Based on CLSM observation, the increase in transfection efficiency was proposed to be due to cellular internalization through caveolae-mediated endocytosis. Additionally, a series of disulfide crosslinked PIC nanoparticles based on oligopeptides and glycopeptides containing both LLys and cysteine residues have been developed by Rice and co-workers.178, 183, 184

Interestingly, disulfide core-crosslinked nanoparticles (CCLNPs) based on polypeptides have been developed by a simple one-step method using a difunctional NCA comonomer. In two separate studies, a difuctional NCA containing a disulfide bond and two NCA rings, i.e., L-cystine-NCA (Scheme 4c), was synthesized.185, 186 Accordingly, disulfide CCLNPs based on PEG and polypeptide was fabricated by one-step ring opening copolymerization of L-cystine-NCA with LPhe-NCA or BLG-NCA using amino-terminated mPEG as macro-initiators. 1H NMR tests indicated that the typical signals of polypeptides were suppressed, suggesting a core–shell structure with a disulfide crosslinked polypeptide core and a PEG shell.185, 186 As shown in Figure8, transmission electron microsopy (TEM) observation revealed that the PEG-poly(LPhe-co-L-cystine) CCLNPs exhibited a uniform spherical morphology with sizes ranging from 50–150 nm, which are smaller than those determined by DLS measurements, likely due to the shrinkage of the nanoparticles during the sample drying process.185 The size of the CCLNPs increased with increasing the overall polypeptide fraction or the content of L-cystine residues (Figure 8A). The CCLNPs remained stable in PBS (pH 7.4) even at an extremely low concentration (1.53 × 10−5 mg mL−1). In contrast, the disulfide crosslinking bonds were cleaved in the presence of 10 mM GSH, rendering an obvious increase in the size of the nanoparticles. Dox-loaded CCLNPs showed a reduction-dependent drug release behavior, as shown in Figure 8B. In PBS without GSH, less than 20% of loaded Dox was released from the CCLNPs at 93.5 h. In contrast, the release was markedly accelerated by addition of GSH, and over 90% Dox was released at 93.5 h in the presence of 10 mM GSH. In addition, an increase in the content of the L-cystine crosslinks or an increase in the overall polypeptide fraction led to a decrease in drug release rate, resulted from an increase in crosslinking density and the formation of a more compact nano-structure. The subsequent CLSM observation revealed effective intracellular delivery of Dox by the Dox-loaded CCLNPs. Strong intracellular fluorescence of Dox was observed within the HeLa cells cultured with the Dox-loaded CCLNPs for 2 h, suggesting intracellular release of Dox from the CCLNPs triggered by GSH-mediated cleavage of the disulfide crosslinks. MTT assays revealed that the polypeptide-based CCLNPs exhibited no significant cytotoxicity and the Dox-loaded CCLNPs showed lower cytotoxicity than free Dox. In addition, a similar GSH-accelerated release of indometacin from a PEG-poly(BLG-co-L-cystine) CCLNP was observed.186 Because the PEG-polypeptide CCLNPs that are prepared through facile one-step ROP of NCAs exhibit an intelligent drug release pattern in response to intracellular reducing environment, it is envisioned that these materials may have potential applications in intracellular drug delivery.

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Figure 8. A) TEM image and the hydrodynamic radii (Rh) of the PEG-poly(LPhe- co -L-cystine) CCLNPs: a) CCLNPs-1 (1/2/19, mPEG/LCys/LPhe (molar ratio)), b) CCLNPs-2 (1/6/23), and c) CCLNPs-3 (1/9/32). B) In vitro release of Dox from the Dox-loaded CCLNPs in PBS at pH 7.4 and 37 °C: traces (a), (b) and (c) refer to the release results of CCLNPs-1, CCLNPs-2 and CCLNPs-3, respectively, without the presence of GSH; traces (d) and (e) represent the results of CCLNPs-3 with the presence of 2.5 mM and 5 mM GSH, respectively; traces (f–h) refer to the results of CCLNPs-1, CCLNPs-2 and CCLNPs-3, respectively, in the presence of 10 mM GSH. Reproduced with permission.185 Copyright 2011, Royal Society of Chemistry.

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In addition, an interlayer-crosslinked micelle has also been reported.187 A PEG-b-PAsp(MEA)-b-PAsp(DIP) triblock copolymer was synthesized by Cu(I)-catalyzed alkyne-azide “click” reaction between azido-terminated PEG-b-poly(N-(2-mercaptoethyl) aspartamide) (PEG-b-PAsp(MEA) and alkyne-functionalized poly(N-(2-(diisopropylamino)ethyl) aspartamide) (PAsp(DIP). At pH 10, the triblock copolymer self-assembled into micelles with a three-layer structure including a PAsp(DIP) core, a PEG outer shell and a PAsp(MEA) intermediated layer (Figure9A-a). Due to the effect of aerial oxidation, disulfide crosslinking bonds were further formed within the PAsp(MEA) interlayer, leading to the formation of interlayer crosslinked micelles. Because of the presence of both tertiary amino groups in the core and disulfide crosslinks in the interlayer, the micelles showed pH- and reduction-sensitive behaviors. At pH 5.0 without adding DTT, an interesting nanocage structure was observed (Figure 9A-b). The nanocages showed a markedly enhanced size than the crosslinked micelles, due to complete dissolution of the inner PAsp(DIP) core that is constrained by the crosslinked interlayer. In addition, with addition of DTT (10 mM) at pH 7.4, highly “swollen” micelles were observed due to cleavage of the disulfide crosslinks and the swelling of partially dehydrated PAsp(DIP) segments (Figure 9A-c). Notably, the nanoparticles disassembled at pH 5.0 with the presence of 10 mM DTT (Figure 9A-d). In vitro drug release measurements revealed that Dox-loaded interlayer crosslinked micelles retained stable in PBS at pH 7.4 without DTT (Figure 9B). In contrast, the release of Dox was significantly accelerated by either reducing the pH to 5.0 or addition of DTT (10 mM), and highest release rate was observed in the presence of dual stimuli. CLSM observation demonstrated that fast accumulation of Dox in the nuclei of Bel-7402 cells was detected for free Dox and the Dox-loaded crosslinked micelles, compared to the distribution of Dox mainly in the cytoplasm for PEG-b-PCL micelles (Figure 9C), indicating an enhanced endosomal/lysosomal escape of Dox for the Dox-loaded crosslinked micelles. Further in vivo studies revealed that the Dox-loaded crosslinked micelles showed higher antitumor efficacy than either free Dox or the Dox-loaded PEG-b-PCL micelles.

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Figure 9. A) TEM images of the nanoassembly at pH values of a) 7.4, b) 5.0, c) 7.4 with addition of DTT, and d) 5.0 with addition of DTT. The interlayer crosslinked micelles shown in (a) were decorated with Au. In TEM measurements, the Au-decorated crosslinked micelles were not stained and other samples were stained with uranyl acetate. The arrows in (b) indicate the “watermark” of staining agent formed as a result of nanocage shrinkage in sample drying. DTT concentration (if added): 10 mM. B) Quantitative Dox release from the dual-sensitive crosslinked micelles (mean ± standard deviation (SD), n = 3). C) Intracellular Dox release and migration into nuclei observed by confocal laser scanning microscopy (CLSM). Bel-7402 cells were incubated (37 °C) for 6 h at a Dox-equivalent dosage of 10 μg per dish. Dox loading contents: 10.5% in the interlayer crosslinked micelles and 5.1% in PEG3k-PCL3k micelles. Nuclei were stained with Hoechst 33342. Reproduced with permission.187

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3.3.2. pH-Sensitive Crosslinked Nanogels

Nanometer-sized polypeptide networks that undergo swelling–deswelling transitions in response to pH change, i.e., pH-sensitive nanogels,170 are interesting for pH-triggered drug-delivery systems. PEG-b-PAsp nanogels were synthesized by crosslinking of the PAsp blocks using 1,6-hexanediamine (HDA) and N,N′-diisopropylcarbodiimide (DIC) as a crosslinker and a coupling agent, respectively.188 A core–shell structure with a crosslinked PAsp hydrogel core and a PEG shell was formed during the crosslinking. The nanogels exhibited a pH-dependent swelling–deswelling transition, as schematically illustrated in Figure10. As the pH increase from 4 to 9, the size of the nanogels increased from below 20 nm to above 40 nm, attributed to the swelling of the polypeptide core caused by the gradual ionization of the Asp residues at pH above its pKa (∼3.9).4 The nanogels showed a constant diameter in 0.15 M PBS (pH 7.4) as the polymer solution was diluted from 5 to 0.2 mg mL−1, compared to the dissociation of PEG-b-PAsp/Ca2+ ion complex micelles at concentrations less than 1 mg mL−1. Dox-loaded nanogels were obtained by mixing Dox with nanogels in deionized water. A relatively high drug loading capacity (26.6 wt%) was obtained, likely due to the electrostatic interactions between oppositely charged Asp residues and Dox. The loading of drug showed no obvious influence on the particle size. It was found that Dox was rapidly released from the nanogels at both pH 5.0 and 7.4, and a slightly faster drug release profile was observed at pH 5.0 than at pH 7.4. It is noteworthy that the nanogels are in a swollen state at 7.4, which may facilitate drug diffusion. The faster release pattern of Dox at acidic pH was assumed to be attributed to an increased solubility of Dox (pKa = 8.25) and reduced interactions between Asp and Dox.

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Figure 10. Schematic illustration of the pH-dependent swelling–deswelling transition of a core-crosslinked PEG-b-PAsp nanogel.

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On the other hand, it was found that the change in the water-solubility of Dox with decreasing pH showed no significant influence on the drug release behavior of a PEG-b-PAsp-b-PLPhe ternary copolymer nanogel.189 PEG-b-PAsp-b-PLPhe triblock copolymers were synthesized via successive ROP of BLAsp-NCA and LPhe-NCA using amino-terminated mPEG as a macro-initiator. Reversible or irreversible nanogels were then prepared by crosslinking the micelles of PEG-b-PAsp-b-PLPhe using an acid-labile ketal-containing crosslinker (Scheme 4d) or a nondegradable crosslinker. The Dox-loaded nanogels containing nondegradable crosslinks exhibited similar drug release profiles at pH 7.4 and 5.0. In contrast, the nanogels crosslinked by acid-cleavable crosslinks showed a markedly enhanced drug release rate as the pH was reduced from 7.4 to 5.0. Hence, an acid-catalyzed hydrolysis of the ketal-containing crosslinks was proposed to be responsible for the accelerated drug release at acidic pH.189 When cultured with MCF-7 cells, the Dox-loaded nanogels with ketal crosslinks resulted in a higher Dox fluorescence intensity within the nuclei than the Dox-loaded nanogels with non-biodegradable crosslinks, indicating an enhanced Dox release from the former nanogels at acidic endosomal/lysosomal environment. Ketal linkages have also been incorporated to the side chains of a polypeptide diblock copolymer, i.e., PEG-b-poly(ketalized serine) (PEG-b-PkSer, Scheme 4e), to fabricate a PEG-b-PLLys analogue with acid-cleavable side chains.190 After crosslinking of PEG-b-PkSer/DNA polyplexes by bis(sulfosuccinimidyl)suberate, crosslinked PIC micelles with both acid-cleavable crosslinks and side chains were obtained. The resulting crosslinked micelles exhibited increased transfection efficiency in the presence of serum than either uncrosslinked PEG-b-PLLys/DNA polyplexes or PEI/DNA polyplexes. CLSM observation revealed an improved dissociation of PEG-b-PkSer and DNA in the cytoplasm compared to the uncrosslinked PEG-b-PLLys/DNA polyplexes, implying an acid-triggered cleavage of the ketal linkages.

Besides the chemical crosslinking methods, photocrosslinking is another commonly used method for in situ crosslinking, for the crosslinking can be performed in a controllable manner with or without the presence of photo-initiators. Chen and co-workers have reported a series of pH-sensitive polypeptide nanogels developed through photo-crosslinking. Photocrosslinkable PEG-b-poly(LGlu-co-γ-cinnamyl L-glutamate) (PEG-b-P(LGlu-co-CLG)) diblock copolymers were synthesized by ROP of BLG-NCA using amino-terminated mPEG as an initiator, followed by grafting cinnamyl alcohol to the LGlu residues.191 PEG-b-P(LGlu-co-CLG) self-assembled into nanoparticles in aqueous solution due to the hydrophobic interactions between CLG segments. Under UV irradiation at 254 nm, the pendant cinnamyl groups within the polypeptide block underwent dimerization, resulting in in situ crosslinking of the polypeptides segments and hence the formation of nanogels. The nanogels showed hydrodynamic radii ranging from 80–135 nm at pH 7.4, depending on the polypeptide block length and the substitution degree of cinnamyl groups. As the pH increased from 4.0 to 7.4, the nanogels exhibited a significant increase in size, due to the swelling of the polypeptide cores caused by gradual ionization of the LGlu residues. MTT assays indicated that both the block copolymers and the nanogels showed no detectable cytotoxicity at concentrations up to 0.1 mg mL−1. Drug-loaded nanogels were prepared by mixing PEG-b-P(LGlu-co-CLG) with a model drug, rifampicin, in aqueous solution, followed by in situ photocrosslinking. The drug-loaded nanogels showed pH-dependent release profiles in vitro. A fast drug release pattern was observed at pH 7.4, compared to only small amount of drug released at pH 4.0, due to the swelling of the nanogels at higher pH.

3.3.3. pH- and Temperature-Sensitive Nanogels

In addition to the nanogels that respond to single stimulus, polypeptide-based nanogels that exhibit swelling–deswelling transitions in response to dual stimuli, such as pH and temperature, have also been investigated for controlled drug delivery. Nanogels based on PNIPAM and PLGlu were synthesized via free radical polymerization of 2-hydroxyethyl methacrylate (HEMA) and PNIPAM grafted PLGlu (PLGlu-g-(HEMA/PNIPAM)) by Chen and co-workers.192 PLGlu-g-(HEMA/PNIPAM) was first prepared by means of conjugating amino-terminated PNIPAM and HEMA to the PLGlu side chains. The nanogels were then obtained by increasing the temperature from 25 °C to 60 °C at pH 8 to form a dispersion of nanoparticles, followed by free radical polymerization of the HEMA residues using ammonium peroxydisulfide (APS) as an initiator. The polymerization of the HEMA residues led to a sharp decrease in the hydrodynamic diameter of the aggregates from 262 nm to about 60 nm, due to the formation of crosslinked nanogels with a more compact structure. The nanogels exhibited pH- and temperature-dependent swelling–deswelling behaviors. At 27 °C, the nanogels showed a decrease in size from ∼70 nm to ∼60 nm as the pH was reduced from 10.0 to 6.0, caused by the gradual protonation of the LGlu segments. Notably, as the pH was further decreased to below 5.5, the particle size increased markedly and large aggregates were detected due to the hydrophobic aggregation of the PLGlu segments. In addition, at pH 7.0, the nanogel size showed a sharp decrease as the temperature was increased to above 36 °C, attributed to hydration-dehydration transitions of the PNIPAM segments. PLGlu/PNIPAM hybrid microgels have also been prepared via free radical copolymerization of HEMA-grafted PLGlu and N-isopropylacylamide (NIPAM) in 0.05 M PBS (pH 7.0) at 60 °C by using APS as an initiator.193 In comparison with the above PLG-g-(HEMA/PNIPAM) nanogels, the microgels exhibited significantly higher size (570 nm) at pH 7.0 and 25 °C. The microgels displayed pH- and temperature-sensitive swelling–deswelling behaviors similar to the nanogels. Because of the dual-stimuli responsive swelling behaviors of the nano-sized and micrometer-sized hydrogel particles, it is envisioned that these materials may be interesting for drug delivery systems, such as oral drug delivery systems.

3.4. Hydrogels

Hydrogels are three-dimentional hydrophilic or amphiphilic polymer networks formed by chemical or physical crosslinking.2, 3 Stimuli-responsive macroscopic hydrogels exhibit volume phase transitions or sol–gel phase transitions in response to changes in environmental conditions.3 Permanent polymer networks can be formed by chemically crosslinking and display volume phase transitions under environmental stimuli, whereas physically crosslinked hydrogels are usually reversible networks and can be reversibly change to an uncrosslinked sol state in respond to external stimuli.3 Therefore, covalently crosslinked hydrogels that undergo volume phase transitions in respond to environmental stimuli have been designed for controlled drug delivery through stimuli-induced swelling–deswelling transitions. On the other hand, in situ forming physically crosslinked hydrogels, such as thermo-gelling hydrogels, have been developed for localized drug delivery.3 Recently, polypeptide-based chemically crosslinked hydrogels and in situ gelling hydrogels have been fabricated and tested for different drug-delivery applications.

3.4.1. Chemically Crosslinked Stimuli-Sensitive Hydrogels

Chemically crosslinked pH-sensitive bulk hydrogels are interesting for pH-triggered drug delivery systems. For example, after being crushed and sieved into microparticles, hydrogels that respond to the pH shift between the stomach (pH ∼ 2) and the small intestine (pH ∼ 7) can be used for oral drug delivery of peptides and proteins.6 Thus, hydrogels based on PLGlu and PAsp containing pendant carboxylic groups have been developed for intestinal drug delivery. Yang and co-workers have synthesized a series of pH-responsive hydrogels based on PLGlu by using diamino-capped PEG as a crosslinker.194 The equilibrium swelling ratios (SR) of the hydrogels exhibited a marked increase with increasing the pH from 4 to 6, due to gradual ionization of the carboxylic groups as well as a secondary conformation transition. An increase in ionic strength led to a decrease in SR at higher pH, owing to the charge screening effect of salt. CD measurements indicated that an ordered secondary structure, i.e., β-sheet, was formed within the hydrogel at pH ≤ 4.5, whereas a random coil structure was observed at higher pH.194 In a subsequent study, PLGlu-based hydrogels were prepared by photoinduced polymerization of PEG-methacrylate substituted PLGlu using PEG dimethacrylate (PEG-DA) as a crosslinker.195 It was found that SR decreased with increasing the crosslinking density, and an increase in the MW of the PEG-DA crosslinker resulted in a decrease in SR at higher pH, probably due to a decrease in the relative charge density. The degradation of the hydrogels was significantly accelerated as the pH was increased. Protein-loaded hydrogels containing insulin, lysozyme or albumin were prepared by a swelling-diffusion method. Initial burst releases were observed for all the protein-loaded hydrogels owing to a rapid diffusion of the proteins located at the hydrogel surface. The in vitro release behavior of the proteins was found to be influenced by the crosslinking density of the hydrogels and the MW of the loaded protein.195 A higher crosslinking density or a higher MW of the protein resulted in a lower release rate at pH 7.4. In addition, pH-sensitive hydrogels based on PLGlu/poly(acrylic acid) (PLGlu/PAA) hybrid systems and PAsp have also been investigated.196–198

pH- and temperature-sensitive hydrogels composed of PLGlu and a thermo-sensitive component have been developed. A series of hybrid hydrogels based on PLGlu and poly(NIPAM-co-HEMA) have been developed by Chen and co-workers.199 The hybrid hydrogels were prepared by coupling PLGlu with poly(NIPAM-co-HEMA) using 1-ethyl-3-(3-dimethylaminopropyl) carbodiimide hydrochloride (EDC•HCl) as a coupling agent. The hybrid hydrogels exhibited swelling behaviors dependent on pH and the hydrogel composition. An increase in the PLGlu content led to an enhanced swelling ratio (SR), a higher pore size and an increased enzymatic degradation rate at pH 7.4. All the hydrogels showed fast swelling-deswelling transitions with varying the pH from 7.0 to 4.0 at 37 °C, indicating potential applications in oral drug delivery. It is noteworthy that the change in temperature showed less influence on the SR of the hydrogels compared to the pH change, even though sharp thermo-induced phase transitions were observed for the hydrogels at around 32 °C. This was believed to be due to the highly hydrophilic nature and electrostatic repulsion of the PLGlu segments at higher pH. A model protein, lysozyme, was loaded in the hydrogels by a swelling-diffusion method. After an initial burst, the in vitro release rate of lysozyme was significantly increased as the pH was increased from 4.0 to 7.0, suggesting that the release of lysozyme was retarded by shrunk hydrogels at acidic pH but accelerated from swollen hydrogels at neutral pH. Very recently, pH- and temperature-sensitive hybrid hydrogels composed of PLGlu and a naturally derived polysaccharide, i.e., hydroxypropylcellulose (HPC), have been prepared by the same group.200, 201 The PLGlu/HPC hybrid hydrogels were synthesized via free-radical copolymerization of HEMA modified PLGlu (PLGlu-g-HEMA) and acrylate substituted HPC.200 Due to the LCST behavior of HPC (LCST ∼ 41 °C),202 the hybrid hydrogels exhibited pH- and temperature-dependent swelling behaviors. The SR of the hydrogels increased markedly with increasing pH from 4.0 to 6.8 at 37 °C, whereas the SR decreased gradually as the temperature was increased from 25 °C to 48 °C at pH 6.8. Interestingly, the HPC content within the hydrogel showed an influence on the SR in an opposite manner at pH ≤ 4.0 and pH ≥ 5.0, respectively. A higher PLGlu content resulted in a higher SR of the swollen hydrogel at pH ≥ 5.0, due to stronger electrostatic repulsion between the PLGlu segments. In contrast, a higher HPC content led to a higher SR of the shrunk hydrogel at pH ≤ 4.0, likely attributed to the fact that a higher HPC content resulted in an enhanced hydrophilicity of the shrunk hydrogel and/or caused a steric hindrance to intermolecular hydrogen bonding.194 Similar to the PLGlu/P(NIPAM-co-HEMA) hydrogels,199 the influence of pH on the swelling behavior of the PLGlu/HPC hydrogels was more marked than that of temperature. As shown in Figure11, a clear swollen hydrogel was observed at pH 6.8 and 25 °C. With increasing temperature to 42 °C at pH 6.8, a turbid hydrogel with slightly reduced SR was observed due to the dehydration of HPC. In contrast, an opaque and shrunk gel with dramatically reduced SR was formed at pH 1.2 and 25 °C, caused by deionization of the PLGlu segments and the formation of strong intermolecular hydrogen bonding. Interestingly, the enzymatic degradation of the hybrid hydrogels was markedly affected by pH and HPC content. The hydrogels with a HPC content of 45 wt% or higher exhibited no obvious degradation in artificial gastric juice (pH 1.2) with the presence of 3.2 mg mL−1 pepsin for 2 h, compared to rapid degradation of the hydrogels with a HPC content of 27 wt% or less. On the other hand, the degradation rate of the hydrogels with the HPC content of 45 wt% or higher was significantly enhanced in artificial intestinal liquid (pH 6.8) with the presence of 10 mg mL−1 pancreatin, due to the swelling of the hydrogels and an increase in enzyme accessibility to the polypeptide backbones. The in vitro release of BSA from the hydrogels in artificial gastric juice (pH 1.2) was suppressed after an initial burst release of BSA located at the hydrogel surface, whereas the release was obviously accelerated in artificial intestinal liquid (pH 6.8).

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Figure 11. pH- and temperature-dependent swelling behaviors of the PLGlu/HPC (45/55, w/w) hybrid hydrogels. Reproduced with permission.200

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3.4.2. In Situ Forming Physically Crosslinked Hydrogels

In situ forming hydrogels or injectable hydrogels can be formed by the in situ formation of physical interactions in response to environmental stimuli,3, 203, 204 or by in situ chemical reactions such as Michael addition reactions,205 enzymatically catalyzed reactions,206, 207 and the reactions of aldehyde with amines.208 In comparison with the permanent polymer networks crosslinked by covalent bonds, physically crosslinked in situ forming hydrogels are reversible networks and can be developed by simply varying the environmental conditions, such as temperature and pH. Due to their unique advantages, such as minimal invasion, no organic solvent, site-specificity, less systematic toxicity and ability to deliver both hydrophobic and hydrophilic drugs, stimuli-sensitve in situ forming hydrogels have received extensive investigation for drug delivery in the past decade.3 For instance, drugs can be mixed with a thermo-sensitive polymer in aqueous solutions at a lower temperature, and drug-loaded hydrogels can be formed in situ after injection of the mixed solutions into body. Recently, stimuli-sensitive in situ gelling hydrogels based on polypeptides have been developed.

Deming and co-workers have developed a series of in situ forming hydrogels by diblock copolypeptides comprising a charged PLLys or PLGlu block and a hydrophobic PLLeu or PLVal block.28, 132 It was found that rigid hydrogels were formed by these diblock copolypeptides at very low polymer concentrations (0.25–2.0wt%). The charged block was found to contribute to gelation at low concentrations. Replacement of positively charged PLLys block with a negatively charged PLGlu block did no affect gel formation. However, addition of salt led to a weaker hydrogel, due to charge screening effect. In addition, the ordered secondary structures, such as α-helix and β-sheet, of the hydrophobic block promoted gel formation. The strength of the hydrogels can be tuned by varying the composition and concentration of the polypeptides. The physically crosslinked polypeptide hydrogels can be deformed and thinned by stress and injected through small-bore cannulae.209 LSCM, ultra SANS, and cryogenic TEM measurements revealed that the hydrogels were composed of an interconnected porous network.210 In vitro cytotoxicity tests indicated that hydrogels containing both PLGlu and PLLys exhibited good cytocompatibility, even though free diblock copolymers containing PLLys are cytotoxic.210 The low cytotoxicity of the hydrogels was proposed to be attributed to the fact that the polypeptide segments were tethered within the hydrogels and not available in solution to cause cytotoxicity. After injection of the PLLys-b-PLLeu hydrogels into the forebrain of mice, the hydrogels exhibited good biocompatibility comparable to physiological saline.209 The hydrogels injected in vivo were found to integrate well with brain tissue, and time-dependent in-growth of blood vessels, certain glia and some nerve fibers into the hydrogels were observed.

In situ thermo-gelling hydrogels are the most widely studied in situ forming systems.3, 203, 204 Biodegradable amphiphilic thermo-gelling hydrogels were firstly developed based on PEG/aliphatic polyesters block copolymers by Kim and co-workers.203 In subsequent studies, biodegradable in situ gelling hydrogels based on different polyesters, polycarbonates and oligo(amidoamine)s were reported.3, 211, 212 The stimuli-induced gelation was believed to be resulted from the synergistic effect of an increase in physical interactions between the hydrophobic segments and partial dehydration of the PEG block.3, 203, 204 Very recently, thermo-gelling systems based on polypeptides have also been reported. Jeong and co-workers synthesized a PEG-b-poly(L-alanine) (PEG-b-PLAla, Scheme5a) diblock copolymer via ROP of L-alanine NCA (LAla-NCA) by using amino-terminated mPEG as a macro-initiator.19 The PEG-b-PLAla aqueous solutions were found to exhibit sol-to-gel transitions with increasing temperature at a polymer concentration of 6–12 wt%. The sol–gel-transition temperatures were in the range of 20–40 °C depending on the polymer concentration, making these materials interesting for in situ forming drug-delivery systems. In contrast, the aqueous solutions of a PEG-b-poly(D,L-alanine) (PEG-b-PAla) diblock copolymer without a secondary conformation only formed hydrogels at higher polymer concentrations (≥16 wt%) and higher temperatures (>60°C). This suggested that the secondary conformation of the PLAla block played an important role in the thermo-induced gelation process. CD and FTIR measurements indicated that an increase in temperature or polymer concentration caused a slight increase in β-sheet content of the PLAla block. Accordingly, the thermo-induced gelation was proposed to be due to an increase in intermicellar aggregation caused by the synergistic effects of the increase in β-sheet content of the polypeptide block and partial dehydration of the PEG block.213 The self-assembled nanostructure and gelation behavior were found to be influenced by the block sequence of the block copolymers.214 The gelation behavior of a PEG-b-PLAla-b-PAla triblock copolymer with a central rigid block was compared with that of a PEG-b-PAla-b-PLAla triblock copolymer. The PEG-b-PAla-b-PLAla aqueous solutions showed sol-to-gel-to-squeezed gel transitions at lower concentrations (4.0–9.0 wt%), compared to only sol-to-gel transitions observed for the PEG-b-PLAla-b-PAla aqueous solutions at higher polymer concentrations, as shown in Figure12a. CD measurements suggested that the PLAla block within both triblock copolymers adopted an α-helical structure at low concentrations (0.01 wt% and 0.025 wt%) and a lower temperature. As the temperature was increased above 40 °C, the secondary conformation of the PLAla block within PEG-b-PAla-b-PLAla exhibited a transition from α-helix to random coil. In contrast, the secondary structure of the PLAla block in PEG-b-PLAla-b-PAla exhibited no significant change within the experimental temperature range (4–50 °C). The α-helix-to-random coil transition was assumed to be responsible for the gel-to-squeezed gel transition of PEG-b-PAla-b-PLAla. As shown in Figure 12b, hydrogels of both triblock copolymers formed at 37 °C displayed highly porous structures, whereas the squeezed gel of PEG-b-PAla-b-PLAla at 60 °C exhibited a collapsed pore structure. In addition, a PEG-b-poly(LAla-co-LPhe) diblock copolymer was also found to show thermo-induced sol-to-gel transitions at concentrations of 3.0–7.0 wt% (Scheme 5b).213 A similar thermo-induced strengthening of β-sheet structure of the polypeptide block was observed for the diblock copolymer. The hydrogels were stable in PBS without enzymes in vitro, whereas the degradation of the hydrogels was significantly accelerated after the hydrogels were subcutaneously injected into rats, due to the presence of enzymes in the subcutaneous layer. In vitro release of insulin from insulin-encapsulated hydrogels showed an initial burst release at the first day, followed by a diffusion-controlled release profile over 16 days. After a single subcutaneous injection of the polymer solution containing insulin into diabetic mice, a significant hypoglycemic effect was maintained over 18 days.

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Scheme 5. Chemical structures of representative thermo-gelling polypeptide-based block copolymers.

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Figure 12. a) Photos of sol (5 °C)/gel (37 °C or 60 °C) states of PEG-b-PLAla-b-PAla (20.0 wt% in water) (A) and sol (5 °C)/gel (37 °C)/squeezed gel (60 °C) states of PEG-b-PAla-b-PLAla (5.0 wt% in water) (B). b) SEM images of gels. Images of the gel of PEG-b-PLAla-b-PAla at 37 °C (A), the gel of PEG-b-PAla-b-PLAla at 37 °C (B-1), and the gel of PEG-b-PAla-b-PLAla at 60 °C (B-2) are compared. The scale bar is 10 μm. Reproduced with permission.214 Copyright 2011, Royal Society of Chemistry.

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In addition, thermo-gelling block copolymers based on poloxamer flanked by two polyalanine blocks were synthesized by ROP of the mixture of LAla-NCA and DL-Alanine NCA (DLAla-NCA) using diamino-terminated poloxamer as a macro-initiator.215 It was found that the sol–gel transition temperature of the copolymer solution was affected by the relative block length and the LAla content within the polypeptide block. Generally, an increase in the polypeptide block length, a decrease in the poloxamer block length, or an increase in the LAla content led to a decrease in the phase transition temperature. The moduli of the in situ fomed hydrogels increased markedly with increasing the polymer concentration.216 Additionally, the polymer concentration showed a marked influence on the biocompatibility of the hydrogels. Hydrogels with polymer concentrations of 7.0 wt% and 10.0 wt% were found to show good biocompatibility and promote the proliferation and differentiation of chondrocytes in vitro and in vivo.216, 217 In contrast, the viability, proliferation and differentiation of chondrocytes cultured within the 15 wt% hydrogels were markedly decreased.216 It was also found that the secondary structure of the polypeptide copolymer showed influence on the nanostructure of the hydrogels as well as the proliferation and differentiation of the encapsulated chondrocytes.217 Block copolymers based on poloxamer end-capped by two poly(LAla-co-LPhe) or poly(LAla-co-LLeu) blocks were found to exhibit similar sol-to-gel transitions.218, 219 As compared with the β-sheet dominant secondary conformations of the copolymers based on PLAla and poly(LAla-co-LPhe), the poly(LAla-co-LLeu)-based copolymer adopted an α-helical structure within the temperature range of 20–50 °C. In addition, in situ gelling PEG/peptide systems containing dipeptides end groups,220 PEG/α-cyclodextrin (α-CD) complexes,221 or an adhesive precursor, i.e., L-3,4-dihydroxylphenylalanine (DOPA),222 have been investigated.

It is noteworthy that some amphiphilic poly(amino acid) derivatives containing both hydrophilic and hydrophobic side groups also show thermo-induced gelation behaviors. Uyama and co-workers synthesized a series of poly(α/β-aspatamide) (PAspAm) derivatives by successive aminolysis of polysuccinimide (PSI) using dodecylamine and amino alcohols.223, 224 The amphiphilic PAspAm derivatives exhibited sol–gel transitions in PBS. The sol-gel transition temperature was found to be influenced by graft composition, the side chain length, polymer concentration and additives. Kim and co-workers developed a series of thermo- and pH-sensitive amphiphilic PAspAm through successive aminolysis of PSI by N-alkylamines and N-isopropylethylenediamine.225, 226 The copolymers showed pH-dependent thermo-sensitivity, and the phase transition temperature was influenced by the alkyl chain length and graft composition. The concentrated copolymer solutions in PBS (pH 7.4) displayed thermo-induced sol–gel transitions in vitro and in vivo. The in vitro release of a hydrophobic model drug, chlorambucil, from the hydrogels followed near zero-order kinetics over 16 days.

4. Stimuli-Sensitive Functionalized Polypeptides

  1. Top of page
  2. Abstract
  3. 1. Introduction
  4. 2. Synthesis of Polypeptides
  5. 3. Stimuli-Sensitive Polypeptide-Based Materials
  6. 4. Stimuli-Sensitive Functionalized Polypeptides
  7. 5. Conclusions and Perspectives
  8. Acknowledgements
  9. Biographical Information

Some polypeptides exhibit responsiveness to environmental stimuli, such as pH and temperature, due to ionization- deionization transitions of the ionizable pendant groups and/or stimuli-induced conformation transitions. However, polypeptide-based materials that exhibit sharp phase-transitions in response to physiologically relevant stimuli, such as a narrow pH change in the region of pH 5.0–7.4 and a temperature change in the range of 10–41 °C, may have advantages in drug and gene delivery applications.5, 15, 18, 19 In addition, the incorporation of bioactive molecules to polypeptide-based biomaterials is required for promoting cell–materials interactions. Hence, the functionalization of polypeptides has attracted increasing attention in recent years. Polypeptides have been functionalized by introduction of various functional groups or stimuli-sensitive moieties to the side chains of the polypeptides.

4.1. Functionalization of Polypeptides

Functionalized polypeptides are mainly obtained through two approaches: ROP of the NCA monomers containing desired functional moieties and postpolymerization modification of polypeptides containing reactive side groups. The first approach leads to the possibility of obtaining functionalized polypeptides via one-step ROP. In contrast, the secondary method facilitates the preparation of polypeptides with different functional moieties. For instance, polypeptides containing oligo(ethylene glycol) (OEG) side chains or sugar residues have been synthesized via ROP of NCAs containing OEG or sugar residues, as shown in Scheme6a.227–232 Recently, more attention has been paid to the approach of postpolymerization modification, such as “click” chemistry,233–237 controlled free-radical polymerization,238–240 aminolysis, and transesterifications.85, 241

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Scheme 6. Synthesis of functionalized polypeptides by the ROP of functionalized NCAs (a), alkyne-azide “click” chemistry (b), thiol-ene “click” chemistry (c), and “grafting from” approaches (d).

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“Click” chemistry has emerged as one of the most widely used methods for modification and functionalization of synthetic polymers, due to its high efficiency and high specificity. Very recently, “clickable” groups such as alkyne, azido and vinyl groups have been incorporated to the side chains of polypeptides. As illustrated in Scheme 6b–d, the functionalization of polypeptides can be effectively achieved by incorporation of reactive groups to the NCA monomers, followed by ROP of the NCAs. In two separate studies, Chen et al. and Hammond et al. synthesized alkyne containing poly(γ-propargyl-L-glutamate) (PPLG) via the ROP of γ-propargyl-L-glutamate NCA (PLG-NCA) by using an alkyl amine as an initiator (Scheme 6b).233, 234 PLG-NCA was prepared by the reaction between LGlu and propargyl alcohol in the presence of sulfuric acid or chlorotrimethylsilane, followed by the reaction with triphosgene. Azido-terminated PEG or azido-substituted sugar molecules were then grafted to the side chains of PPLG via Cu(I)-catalyzed Huisgen azide-alkyne 1,3-dipolar cycloaddition reaction. The grafting efficiencies were found to be near 100%. In addition, “clickable” azido-functionalized PLGlu were also developed by Zhang and co-workers.235, 242 Poly(γ-3-azidopropanyl-L-glutamate) (PAPLG) was obtained by the reaction between sodium azide and Poly(γ-3-chloropropanyl-L-glutamate) (PCPLG), which was synthesized via the ROP of CPLG-NCA by using hexamethyldisilazane (HMDS) as an initiator. Similarly, alkyne-substituted sugar moieties and alkyne-terminated PEG-b-PLA were then grafted to the polypeptide side chains via azide-alkyne ‘click’ chemistry. It is noteworthy that the chloride-substituted side groups of polypeptides can also be used as initiators for atom transfer radical polymerization (ATRP). A highly efficient “grafting from” method by using the chloride-functionalized side groups as initiators for the modification of polypeptides has been developed by Chen and co-workers (Scheme 6d).238–240

Thiol-ene “click” chemistry is another versatile and efficient method for the functionalizaiton of polymers and biomolecules. Polypeptides containing vinyl side groups have been developed for the functionalization of side chains. Schlaad and co-workers developed a poly(D,L-allylglycine) (PAGly) via ROP of D,L-allylglycine NCA.232 Glycopolypeptides were obtained via the thiol-ene “click” reaction between PAGly and 1-thio-β-D-glucopyranose under UV irradiation. A similar allyl containing polypeptide, i.e., poly(γ-allyl-L-glutamate) (PALG), was reported by Zhang and co-worker.243 PALG was synthesized via ROP of γ-allyl-L-glutamate NCA (ALG-NCA) (Scheme 6c). In addition, multi-functional polypeptides with both allyl and azido side groups were obtained via successive ROP of ALG-NCA and CPLG-NCA, followed by replacement of chloride with azide. As a result, functional moieties can be further introduced to the polypeptides by either thiol-ene or azide-alkyne “click” reaction. Cheng and co-workers synthesized a vinyl containing poly(γ-4-vinylbenzyl-L-glutamate) (PVBLG) via ROP of γ-4-vinylbenzyl-L-glutamate NCA (VBLG-NCA) by using HMDA as an initiator.237, 244, 245 In addition to the “click” reactions with thiol-substituted molecules,237 the vinyl groups of PVBLG can also be converted to other functional groups, such as alcohol, aldehyde, carboxylic acid, or bishydroxyl groups.244, 245

Other popular approaches for modification of polypeptides include aminolysis and transesterifications.85, 241 Poly(aspartamide)s containing various functional groups have been obtianed via aminolysis of PBLA or PSI.85, 223, 224 PLGlu derivatives have been fabricated via aminolysis of PBLG.246, 247 In addition, PBLG modified with different reactive groups, such as azido group, propargyl group, allyl group and chloro-substituted group, have also been developed via ester-exchange reactions, and the degree of substitution of the functional groups was found to be dependent on the feed ratio of alcohol to PBLG.241

4.2. Thermo-Sensitive Functionalized Polypeptides

Although some polypeptides show thermo-responsiveness, such as thermo-induced change in secondary conformation,153 intelligent polypeptides that exhibit sharp thermo-dependent phase-transitions at temperatures near the physiological condition (such as 10–42 °C) may have advantages in practical drug delivery applications.18, 19 Consequently, polypeptides functionalized by thermo-sensitive moieties have been developed. Oligo(ethylene glycol) (OEG) grafted polymers, e.g., MEGx-grafted polymethacrylate (denoted as P(MEGxMA), where the subscript numbers represent the number of EG residues within each oligomer segment), have been shown to be thermosensitive, and their LCST can be adjusted by varying the MEGx side chain length. Polypeptides with various MEGx side chains have also been synthesized recently. Chen and co-workers have synthesized a seres of MEGx-grafted PLGlu (PPLG-g-MEGx) via the “click” reaction between poly(γ-propargyl-L-glutamate) (PPLG) and azido-substituted MEGx (Scheme 6b).248 PPLG-g-MEGx exhibited sharp thermal phase transitions with the LCST depending on the polypeptide backbone length, MEGx side-chain length, polymer concentration and salt concentration. An increase in the MEGx side-chain length or a decrease in the polypeptide backbone length led to an increase in LCST. As a result, the LCST of the polypeptide derivative could be tuned from 22 °C to 74 °C, as shown in Figure13. CD measurements suggested that the polypeptide derivatives adopted an α-helical structure. MTT assays indicated that the polypeptide derivatives exhibited no detectable cytotoxicity. In vitro degradation of the polypeptides was observed in the presence of proteinase K. In a separate investigation by Li and co-workers, poly(γ-MEGx-L-glutamate) (P(MEGxLG)) was obtained by synthesis of NCAs containing MEGx, followed by the ROP of the NCA monomers (Scheme 6a).228 P(MEGxLG) with mixed MEGx side chains, i.e., MEG2 and MEG3, were fabricated via ROP of the mixture of MEG2-L-glutamate NCA and MEG3-L-glutamate NCA. It was found that the LCST could be tuned from 32 °C to 57 °C by adjusting the NCA feed ratio. Temperature-dependent CD measurements indicated that the secondary structures of the polypeptides contributed to the LCST behaviors, even though the secondary structures were not disrupted at temperatures above the LCST. It is noteworthy that P(MEG2LG) predominantly existed as a β-strand structure and eventually lost its LCST behavior, compared to an α-helical conformation of PPLG-g-MEG2.248 The difference is likely due to the effect of the triazole ring within the side chains of PPLG-g-MEG2.

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Figure 13. Thermal phase transitions of 10 mg mL−1 aqueous solutions of oligo(ethylene glycol) functionalized poly(L-glutamate)s (denoted as PPLGn-g-MEOx, where the subscript numbers refer to the average number of the repeat units within each segment). Reproduced with permission.248 Copyright 2011, Royal Society of Chemistry.

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Thermo-sensitive PLGlu derivatives have also been developed via a “grafting from” strategy by Chen and co-workers.238, 239 PLGlu-g-P(MEGxMA) was synthesized via ATRP of MEGx methacrylate (MEGxMA) by using poly(γ-2-chloroethyl-L-glutamate) (PCELG) as a macro-initiator (Scheme 6d). The LCST of the graft copolymers was influenced by both the MEGx chain length and the degree of polymerization of P(MEGxMA). CD measurements indicated that the copolymers adopted an α-helical conformation in aqueous solutions, and the secondary conformation was not disrupted with the increase in temperature.

It is noteworthy that some poly(amino acid) derivatives functionalized by amphiphilic pendant groups exhibit LCST behaviors. Kobayashi and co-workers have fabricated poly(N-substituted α/β-aspartamide)s by aminolysis of PSI using a mixture of 5-aminopentanol and 6-aminohexanol.249 The LCST of the poly(N-substituted α/β-asparagine)s increased from 23 to 44 °C as the content of pentanol side groups increased from 50% to 80%. A range of pH- and thermo-responsive poly(N-substituted α/β-aspartamide) derivatives have been obtained by incorporation of both alkylalcohol and alkylamino pendant groups.250, 251 In addition, thermo-responsive poly(γ-glutamic acid) derivatives functionalized by amino alcohols or hydrophobic side groups have also been reported.252–254 Because of the presence of carboxylic side groups, the LCST of the poly(γ-glutamic acid) derivatives showed dependence on pH and salt concentration.

4.3. pH-sensitive Functionalized Polypeptides

Smart polymers that sharply respond to a narrow pH change within the pH region of pH 5.0–7.4 are interesting for intracellular and anticancer drug-delivery systems. Therefore, the development of pH-sensitive polypeptide derivatives capable of responding to a narrow pH change near the physiological pH has received increasing investigation. Hammond and co-workers synthesized a series of PLGlu derivatives containing various amine side groups, including primary, secondary and tertiary amine groups, via the “click” reactions between poly(γ-propargyl-L-glutamate) (PPLG) and azido-substituted alkylamines (Scheme 6b).255, 256 It was found that all the polypeptide derivatives containing primary, secondary and tertiary amine pendant groups showed proton buffering capacity in the pH range of 5.0–7.4. The primary and secondary amine-functionalized polypeptides displayed broad proton buffering behavior in the pH range of 5.5–10 with a midpoint at 7.25. The buffering behavior of the primary and secondary amine-functionalized PPLG at pH lower than the pKa (9–11) of the primary and secondary amines was believed to be due to segmental charge repulsion. For the tertiary amine-functionalized PPLG, the buffering behavior was found to be influenced by the hydrophobicity of the pendant group. The diisopropylamine-functionalized PPLG displayed a sharper buffering transition and a lower pKa compared to the diethylamine- and dimethylamine-functionalized PPLG, as listed in Table 1. It is notable that the triazole ring showed slight effect on the buffering behavior at pH 3–4. CD measurements revealed that the polypeptides adopted an α-helical conformation at high pH values, whereas random coil structure was detected with decreasing pH to below the pKa. The pH-dependent water solubility of the amino-functionalized polypeptides was tested. The primary amine-, secondary amine-, and dimethylamine-substituted PPLG show no detectable phase transition within the experimental pH range. This is different from the solution behavior of amino-functionalized poly(2-hydroxyethyl methacrylate)s (PHEMA) that are also synthesized by grafting the amino groups to PHEMA via “click” chemistry.257 Sharp pH-induced phase-transitions were observed for the primary amine- and dimethylamine-functionalized PHEMA at pH 12.3 and 9.7, respectively. On the other hand, the diethylamine- and diisopropylamine-functionalized PPLG exhibited sharp pH-dependent phase-transitions, as shown in Figure14.255 The phase transition pH decreased with the increase in the hydrophobicity of the tertiary amine group or the DP of the polypeptide. The phase transition pH can also be tuned by incorporation of both tertiary amines to the polypeptides at a given ratio. The diblock copolymers comprising PEG and a tertiary amine-functionalized PPLG showed pH-dependent micellization–demicellization transitions at pH near the physiological condition, making these copolymers interesting for pH-triggered drug release systems. In addition, the hydrolysis of the ester bonds in the side chains was found to be accelerated with increasing pH, but it was markedly reduced by the incorporation of the PEG block, due to the formation of a polypeptide core encapsulated by PEG at higher pH.

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Figure 14. Transmission as a function of pH for all diethylamine and diisopropylamine functionalized polymers. Diethylamine is abbreviated DE and diisopropylamine is abbreviated DI. Reproduced with permission.255 Copyright 2011, Royal Society of Chemistry.

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Amino-functionalized polypeptides have also been developed via “grafting from” approach and aminolysis. Chen and co-workers have synthesized an oligo(2-aminoethyl methacrylate) (OAEMA) grafted PLGlu (denoted as PLGlu-g-OAEMA) by ATRP of 2-aminoethyl methacrylate (AEMA) using PCELG as an initiator (Scheme 6d).240 Similar to the primary amine-functionalized PPLG,255 PLGlu-g-OAEMA exhibited a pKa (7.3) much lower than that of primary amines, due to the segmental Coulombic repulsion. In vitro cytotoxicity assay indicated that PLGlu-g-OAEMA showed a lower cytotoxicity than PEI (MW 25,000), and gel retardation assay displayed that PLGlu-g-OAEMA effectively condensed DNA at a polymer/DNA weight ratio of 0.3 or higher. In addition, typical pH-responsive polypeptide derivatives prepared by aminolysis include diamine-functionalized poly(α/β-aspartamide)s and poly(L-glutamine)s containing two types of amino side groups.85, 246 Due to the presence of two types of amins in the side chains, the polypeptide derivatives exhibited a combined buffering behavior, leading to gene delivery systems capable of both forming complexes with DNAs and promoting endosomal escape.

5. Conclusions and Perspectives

  1. Top of page
  2. Abstract
  3. 1. Introduction
  4. 2. Synthesis of Polypeptides
  5. 3. Stimuli-Sensitive Polypeptide-Based Materials
  6. 4. Stimuli-Sensitive Functionalized Polypeptides
  7. 5. Conclusions and Perspectives
  8. Acknowledgements
  9. Biographical Information

In this Review, we have presented a summary of the recent stimuli-sensitive synthetic polypeptide-based materials that have been designed and tested for drug- and gene-delivery applications. Stimuli-sensitive synthetic polypeptides exhibit unique secondary conformation transitions in response to external stimuli, such as pH and temperature. The unique stimuli-responsive secondary conformations of polypeptides lead to interesting self-assembly behaviors. In addition, the presence of functional groups within the polypeptide side chain facilitates the incorporation of different stimuli-sensitive moieties, such as acid-labile linkers and reduction-sensitive bonds. Besides the above properties, the biodegradability, nontoxicity and the ability to form complexes with biopharmaceuticals and ionic molecules of polypeptides make these materials promising candidates for drug- and gene-delivery applications. Polypeptide-based materials including micelles, vesicles, nanogels, and hydrogels have been developed based on the polymers with different structures. Structure and/or phase transitions of these materials, such as micellization–demicellization transitions, crosslinking–decrosslinking reactions, swelling–deswelling transitions, and sol–gel transitions, can be triggered by varying environmental conditions, such as pH, temperature, redox environment, and dual stimuli. Accordingly, these materials have been evaluated for different drug-delivery applications, including anticancer drug and gene delivery, oral delivery of proteins, and localized drug delivery.

The basic reqirements for applying these materials in drug and gene delivery include efficiency, specificity, and safety.74 To achieve desirable therapeutical objectives, drug-delivery systems with multifunctionalities and the ability to respond to multiple physiologically relevant stimuli have received increasing attention. For instance, disulfide-crosslinked pH-sensitive PEG-polypeptide nanocarriers have advantages including high stability in extracellular environment, improved endosomal escape and enhanced intracellular delivery. Gene carriers with both DNA condensation ability and buffering capacity at endosomal pH can improve endosomal disruption and transfection efficiency. pH- and thermo-sensitive hydrogels can realize drug loading and release by adjusting either pH or temperature. Besides the ability to respond to physical and chemical stimuli, incorporation of biofunctionality to drug-delivery systems plays a crucial role in promoting cell–material interactions. Owing to the presence of functional and/or ionizable groups in some polypeptides, multi-functional polypeptide-based materials can be readily obtained by incorporation of different functional moieties to the polypeptides. On the other hand, it is worth mentioning that even though polypeptides and their derivatives containing not more than two kinds of amino acid residues exhibit no or less immunogenicity, some synthetic random copolypeptides comprising three (or more) kinds of amino acid may be immunogenic in vivo.258–260 Therefore, the amino acid composition of synthetic polypeptides needs to be considered during the molecular design of polypeptide-based materials for in vivo applications. Additionally, incorporation of hydrophilic and biocompatible segments, e.g., PEG, can markedly reduce the immunogenicity and significantly improve the biocompatibility of the synthetic polymeric materials.

Acknowledgements

  1. Top of page
  2. Abstract
  3. 1. Introduction
  4. 2. Synthesis of Polypeptides
  5. 3. Stimuli-Sensitive Polypeptide-Based Materials
  6. 4. Stimuli-Sensitive Functionalized Polypeptides
  7. 5. Conclusions and Perspectives
  8. Acknowledgements
  9. Biographical Information

The authors are grateful for the financial support from the National Natural Science Foundation of China (projects 51003103, 21174142, 50973108, 20904053 and 21074129), the Ministry of Science and technology of China (International cooperation and communication program 2010DFB50890), and the Scientific and Technological Development Projects of Jilin Province (201101082 & 20110332).

Biographical Information

  1. Top of page
  2. Abstract
  3. 1. Introduction
  4. 2. Synthesis of Polypeptides
  5. 3. Stimuli-Sensitive Polypeptide-Based Materials
  6. 4. Stimuli-Sensitive Functionalized Polypeptides
  7. 5. Conclusions and Perspectives
  8. Acknowledgements
  9. Biographical Information
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Xuesi Chen received his Ph. D. at Waseda University, Japan, in 1997, and completed his post-doctoral fellowship at the University of Pennsylvania, USA, in 1999. He has been a full Professor at Changchun Institute of Applied Chemistry, Chinese Academy of Sciences, since 1999. His research interests focus on the development and biomedical applications of biodegradable polymers and intelligent biomaterials, mainly based on polyesters, polypeptides, polycarbonates, and their copolymers.