Responsive polymers for medical diagnostics

Divambal Appavooab, Sung Young Parkb and Lei Zhai*a
aNanoScience Technology Center, Department of Materials Science and Engineering, Department of Chemistry, University of Central Florida, FL 32826, USA. E-mail:
bDepartment of Chemical and Biological Engineering, Korea National University of Transportation, Chungju 380-702, Republic of Korea

Received 10th February 2020 , Accepted 10th May 2020

First published on 19th May 2020

Stimulus-responsive polymers have been used in improving the efficacy of medical diagnostics through different approaches including enhancing the contrast in imaging techniques and promoting the molecular recognition in diagnostic assays. This review overviews the mechanisms of stimulus-responsive polymers in response to external stimuli including temperature, pH, ion, light, etc. The applications of responsive polymers in magnetic resonance imaging, capture and purification of biomolecules through protein–ligand recognition and lab-on-a-chip technology are discussed.

1. Introduction

Medical diagnostics is the process (e.g. diagnostic assay and diagnostic image) used to determine which disease or condition explains a patient's symptoms and signs. It is a complicated and crucial stage in biomedicine because early detection of diseases is of the utmost importance as it increases the chances of curing. For example, selecting an appropriate diagnostic assay for infectious diseases is determined by multiple parameters including clinical presentation and endemic pathogens known to circulate within a specific geographical region. Rapid point-of-care polymerase chain reaction (PCR)1 and lateral flow immunoassays2 and laboratory-based antigen capture enzyme-linked immunosorbent assays (ELISAs)3,4 can generate a clinically actionable diagnosis of patients in that region. These approaches require efficient separation and purification of the biomolecules of interest to obtain the information in a timely manner. The diagnostic imaging describes a variety of non-invasive methods such as computed tomography (CT), magnetic resonance imaging (MRI), ultrasound, etc. that enable the inside view of the body to determine the cause of an injury or an illness. The image resolution and contrast between normal tissues and pathological tissues are desirable for accurate and early detection.5 Detecting diseases in a fast, sensitive, accurate and cost effective manner has been the Holy Grail for medical research. The rise and fall of Elizabeth Holme's Theranos reflected the need, the opportunity and the challenge of developing an efficient diagnosis method.

Responsive polymers can change their chemical and physical properties in response to external physical (temperature, light, solvents, electric fields, magnetic fields, and ionic strengths), chemical (pH, specific ions or chemical agents) and biological (enzyme substrates, affinity ligands and other biological agents) stimuli.6–22 The advancement in polymerization techniques and the understanding of structure–property relationships has granted stimulus-responsive polymers many applications in the biomedical field including biosensors, tissue engineering, drug discovery, drug delivery, and medical diagnostics, as evidenced by the number of reviews reported in this field.20–28

Stimulus-responsive polymers have been used in improving the efficacy of medical diagnostics through different approaches including enhancing the contrast in imaging techniques and promoting molecular recognition. For example, contrast agents functionalized by pH-responsive polymers in magnetic resonance imaging (MRI) can be selectively activated in pathological tissues, providing upgraded signals and more efficient disease detection.15,29,30 Molecular recognition is important in identifying diseases because it allows the selective isolation and detection of biomolecules of interest based on interactions/affinity. The conjugation of stimulus-responsive polymers with biomolecules has facilitated molecular recognition-based separation and purification of biomolecules.31

Another emerging field in biomedicine is lab-on-a-chip technologies, which involve the miniaturization of processes that take place on an integrated system. Stimulus-responsive polymers on lab-on-a-chip devices have contributed significantly to the advantages of miniaturization, such as reduction in cost, time, reagents and samples required, sensitivity and parallelization.32–35 The applications of lab-on-a-chip devices based on stimulus-responsive polymers include sensing,36 drug delivery,37–40 diagnosis,23,41,42 etc.

Many polymer systems developed for their biomedical applications such as drug delivery and diagnosis have been nicely discussed in recent review articles.15,24,31,43–47 For example, Peppas and coworkers have overviewed the strategies of making analyte-responsive hydrogels for sensing and controlled release based on competitive molecular interactions or physicochemical changes.48 The synthesis of nanoscale architectures of core-brush and core–shell nanoparticles, and hybrid nanogels have been reviewed for therapy and diagnosis applications.49 Some review articles present the design, synthesis and applications of responsive polymers in specific fields including efficient probing of circulating tumor cells,50 treating and diagnosing cardiovascular diseases,44 enhancing targeted MRI, in vivo photodynamic therapy, etc.51

This article provides an introductory overview of the mechanisms of stimulus-responsive polymers in medical diagnosis as they respond to external stimuli including temperature, pH, ion concentration, protein interactions, light irradiations and magnetic fields. The application of responsive polymers in diagnosis such as imaging and molecular recognition is discussed based on their responsive mechanism. The mechanisms and applications of stimulus-responsive polymers in the lab-on-a-chip technology focus on the controllable actuators and separation/purification of biomolecules. The conclusions and perspectives summarize the achievements, challenges and future of using the responsive polymers in medical diagnosis, and deliberate the potential of polymer-based microstructures such as microneedles in diagnosis.

2. Responsive mechanisms

2.1. Temperature

Temperature-responsive polymers have found applications in different areas including controlled drug release, flow control and smart coatings.12,16,52–55 The temperature-responsive capability of polymers such as poly(N-isopropylacrylamide) (PNIPAM) and poly(ethylene oxide)-b-poly(propylene oxide)-b-poly(ethylene oxide) (PEO–PPO–PEO)56 is attributed to the hydrophobic interactions or hydrogen bonds among polymer chains and hydrophilic interactions between polymer chains and water molecules. The temperature around which the hydrophobic–hydrophilic transition takes place is referred to as the critical solution temperature and exists as the upper critical solution temperature (UCST) or lower critical solution temperature (LCST). A polymer that possesses UCST exists in one phase at T > UCST and undergoes phase separation at T < UCST. In contrast, a polymer possessing LCST undergoes phase separation as the solution temperature is raised above LCST.

PNIPAM is a common temperature-responsive polymer that demonstrates a negative temperature response.57–60 PNIPAM is soluble in water at temperatures below the LCST (33 °C) where the polymer chains form hydrogen bonds with the water molecules. The polymer chains undergo a coil–globule transition as the surrounding temperature approaches the LCST, and become insoluble in water when the temperature reaches LCST. The phase transition of the PNIPAM is associated with the reversible hydrogen bonding between the hydrophilic amide groups of the polymer and water molecules. When the temperature of the medium is raised, the hydrogen bonds among the amide groups and the water molecules break followed by the formation of intramolecular hydrogen bonds among the amide groups. The exposure of the propyl moieties increases the polymer hydrophobicity accompanied by the release of trapped water and a quick collapse of the polymer from its solution.61–64 The properties of temperature-responsive polymers such as LCST can be changed by modifying the polymer via copolymerization with other monomers, or via addition of other components to the polymer solution.65

A triblock pluronic copolymer, PEO-b-PPO-b-PEO, has demonstrated an interesting temperature-responsive behavior attributed to the interactions of hydrophobic PPO and hydrophilic PEO blocks. A small change in the temperature has resulted in a drastic shift in the critical micelle concentration (CMC) of the copolymer. The PEO chains interacted with water molecules and adopted an extended conformation below micellization temperature (25 °C). In contrast, when the temperature was increased above micellization temperature (37 °C), the interactions between the PEO and water molecules were reduced, dehydrating the polymer chains. Consequently, the interactions between PPO/PPO chains and PEO/PEO chains increased, triggering the shrinkage of the copolymer network.66–69

2.2. pH

pH-Responsive polymers are extensively used in biomedical applications because the pH in pathological tissues is different from the normal physiological pH.13,17,70–72 For instance, the pH in tumor tissues (∼6.6) is lower than that in normal tissues (7.4).73 This pH difference causes the structural change of pH-responsive polymers, granting the capability of detecting diseases.56

Many polymers with amine groups undergo structural changes when the amine groups are protonated. For example, 2-(N,N-diethylamino) ethyl methacrylate (EAMA) in poly(2-[N,N-diethylamino]ethyl methacrylate) (PEAMA) underwent a change from hydrophobic to hydrophilic upon the decrease of the environmental pH. The protonation of the tertiary amino groups of PEAMA in acidic conditions triggered the swelling of the polymer.74 When copolymerized with polyethylene glycol (PEG), the PEAMA core was tethered to PEG chains, generating PEGylated nanogels with pH-responsive properties. The nanogel swelled at low pH because of the protonation and solvation of the PEAMA core and the increase of the osmotic pressure. Furthermore, a fluorinated derivative of PEAMA, poly(2,2,2-trifluoroethyl methacrylate) (PTFEMA) exhibited pH-responsivity. When PTFEMA was tethered to PEG, the fluorinated PEGylated nanogel size increased from 60 nm to 75 nm when the environment pH decreased from 7.0 to 6.5.75–77

Poly(methyl methacrylate-4-vinylpyridine) (PMMAVP) increases its hydrophilicity through the protonation of the pyridine moiety when the solution pH decreases. The copolymerization of hydrophilic PEG and hydrophobic PMMAVP has generated polymer vesicles that respond to the pH changes of the solution. Gold nanoparticles were attached to pyridine groups to monitor the vesicle structure changes through localized surface plasmon resonance (LSRP). When the pH of the solution was lowered from 7.4 to 5.0, the LSPR peak exhibited a blue-shift, indicating an increase in the interparticle spacing. This change suggested the dissolution of the PMMAVP brushes and the disruption of the vesicles caused by the protonation of vinylpyridine.78,79

Similarly, amine groups in polyethylenimine (PEI) are readily protonated when the environment pH decreases, making PEI a “proton sponge”. For example, the –NH2 groups in PEI generated electrostatic repulsions that caused an osmotic swelling of PEI inside the acidic lysosomal intracellular compartments (pH = 5) and a subsequent rupture of the lysosome. The hydrophilic–hydrophobic balance of the system could be adjusted by chemically modifying the polymer chain, which would in turn change the “proton sponge capacity” of the PEI.68,80,81

Poly(2-dimethylaminoethyl methacrylate) (PDMAEMA) had a hydrophilic–hydrophobic transition as the surrounding pH changed. The tertiary amine groups of DMAEMA were protonated at pH 6.0, leading to an increased charge density. This increased internal osmosis causes the swelling of PDMAEMA chains. In contrast, the ammonium groups were deprotonated when the pH increased to 8.0, causing the collapse of the polymer chains and the formation of intramolecular hydrogen bonds. The polymer chains became more hydrophobic and shrank to a smaller hydrodynamic size.82,83

Following the same ionization mechanism, polymers with carboxylic acid groups change their configuration when the degree of ionization of acid groups is changed by the pH, generating different concentrations of –COOH groups and –COO ionic groups at different pH values. Copolymerizing temperature-responsive PNIPAM and pH-responsive acrylic acid (AAc) has produced polymers with dual-responsive properties. The responsive ability of the copolymer could be tuned by changing the AAc[thin space (1/6-em)]:[thin space (1/6-em)]NIPAM ratio. For instance, increasing the percentage of AAc in the copolymer increased the concentration of carboxylate groups, making the copolymer more hydrophilic at pH 7.4 and increasing the LCST of the copolymer. It was demonstrated that increasing AAc concentration from 2% to 6% resulted in an increase of LCST from 33 °C to 55 °C. However, LCST did not change obviously with AAc concentration change at pH 4.0. At low pH, small amounts of AAc were ionized and did not make significant impact on the hydrophilicity of the copolymer.84,85

2.3. Glucose

Phenylboronic acid (PBA) has been used to functionalize polymers to provide binding sites for glucose. PBA produced a boronate ion when it reacted with hydroxyl groups. The tetrahedral boronate bound to glucose to form a stable conjugate. Increasing the concentration of glucose shifted the equilibrium to produce more negatively charged boronate–glucose conjugates (Fig. 1a). As discussed previously, LCST of temperature-responsive copolymers containing PNIPAM depends on the concentration of ionized groups in the polymers.84,85 Functionalizing PNIPAM with 3-aminophenylboronic acid has generated microgel particles whose volume phase transition temperature (VPTT) changed with glucose concentration.86–88 Fig. 1b shows the changes in the microgel particle size (i.e. hydrodynamic diameter) and VPTT with temperature (from 22 to 37 °C) and glucose concentration (from 0 to 3 mg mL−1). The glucose in the solution generated more ionized boronate–glucose conjugates in the microgel, leading to an increase of VPTT. The effect of glucose concentrations on the swelling-collapsing states of the particles was observed in the temperature range of 25 °C to 32 °C (onset-VPTT in the absence of glucose = 25 °C and offset-VPTT in the presence of glucose = 32 °C). For example, the volume change of the microgels started at 25 °C and stopped at 30 °C in a solution free of glucose. In contrast, the microgels in a solution of 3.0 mg mL−1 glucose still underwent volume change below 33 °C. In addition, the ionized boronate–glucose conjugates in microgels increased the degree of swelling due to the increase of osmotic pressure. At temperatures below VPTT, the ionization equilibrium between PBA and glucose determined the swelling of the microgel. The volume change with temperature increased with the glucose concentration. An eight-fold increase in volume of the microgel was observed when the concentration of glucose was increased from 0 mg mL−1 to 2 mg mL−1 at 30 °C. But this ionization-induced swelling was suppressed by the hydrophobic interactions at the temperature above VPTT because the microgels collapsed.
image file: d0tb00366b-f1.tif
Fig. 1 (a) A chemical route for reversible binding of glucose to (alkylamido)phenylboronic acids. (b) Hydrodynamic diameter changes with temperature in the presence of physiologically relevant glucose concentrations at pH 9. The hash marks along the x-axis illustrate the systematic changes in the midpoint VPTT value of the microgel at each plotted glucose concentration. Reproduced from ref. 88, with permission from the American Chemical Society, copyright 2007.

2.4. Ions

Ions in aqueous solutions change polymer structures through the disruption of hydrogen bonds between water molecules and polymers. One example of the ion-responsive polymer is copolymer oligo(ethylene glycol) methyl ether methacrylate (OEGMA) and 2,2,2-trifluoroethyl acrylate (TFEA) (poly(OEGMA-co-TFEA)). The copolymer exhibited thermal- and ionic-responsivities. The thermal sensitivity resulted from the change of H-bonding between water molecules and OEGMA with temperature changes. The copolymer changed their structures upon exposure to anions, such as Cl and SO42−, because of the disruption of hydrogen bonds between water and the copolymer by these ions. When salt was introduced to an aqueous copolymer solution, the ions interacted with the water molecules around the hydrophilic OEGMA chains, reducing the hydrogen bonding between water molecules and OEGMA chains. With the contribution from the hydrophobic TFEA chains, ions can cause a phase separation of this copolymer without a change in temperature.89–91

2.5. Light

Light is another parameter that can be used to trigger responses from the photosensitive polymers. The advantages of this stimulus are the broad spectrum and the high accuracy with which light can be applied.92 Polymers respond to light usually through three routes, switch of polymer hydrophilicity and conformation through crosslinking93 or a breakage of functional groups like nitrobenzyl groups,15,94–97 polymer conformation change attributed to light-sensitive moieties such as azobenzene56,98 and spiropyran,99 and generating fluorescence upon irradiation.100–102

Poly(2-nitrobenzyl acrylate) (PNBA) is a hydrophobic photoresponsive polymer that undergoes dissociation to hydrophilic poly(acrylic acid) (PAA) upon irradiation. The degradation of the PNBA takes place by the photolysis of ester cleavage of the o-nitrobenzyl protective group by long wavelength UV irradiation. The photolysis generates acid groups, which are responsible for shifting the hydrophilic/hydrophobic balance.94–96

UV irradiation has caused the hydrophobic 2-nitrobenzyloxycarbonylaminoethyl methacrylate (NBOC) groups in an amphiphilic poly(ethylene oxide)-b-poly(2-((((2-nitrobenzyl)oxy)carbonyl)amino)ethyl methacrylate) (PEO-b-PNBOC) copolymer to release 2-nitrobenzaldehyde, making the polymer more hydrophilic.103

A hydrogel containing PNIPAM, acrylic acid and spiropyran was reported to change the size upon light irradiation.99 In this system, spiropyran was a light responsive moiety and acrylic acid was used to shift the equilibrium of spiropyran toward the polar protonated merocyanine. Upon exposure to blue or white light-emitting diode (LED) light, the polar merocyanine isomerized to the apolar spiropyran and released the bonded water, causing the shrinkage of the hydrogel (Fig. 8a).

3. Responsive polymer applications

3.1. Magnetic resonance imaging

Magnetic resonance imaging (MRI), since first used in 1977, has become a life-saving technology that allows a non-invasive detection of diseases.104 This imaging technology is based on the analysis of atom behavior in a magnetic field by monitoring the relaxation rates of nuclear spins. Nuclei used for living organisms include 1H and 19F nuclei, with 1H being the most widely used nucleus. However, MRI is not sensitive enough for use in diagnosing diseases since it does not provide sufficient contrast between pathological and normal tissues. Contrast agents were therefore developed to boost the sensitivity of MRI. The role of the contrast agent is to enhance the MRI signal by lowering the longitudinal (T1) or transverse (T2) relaxation times of the nuclei. Relaxivity, which is the change in relaxation rate of water protons, is determined by several factors, such as rotational correlation time, hydration number and the mean residence lifetime of bound water molecules. Gadolinium ion complexes (Gd3+) are the most common T1 contrast agents because they can change the relaxation time of nearby molecules.105

MRI contrast agents need to have the appropriate size to avoid rapid clearing from the body. Conjugating the contrast agents with polymers not only improves the retention time of the contrast agents, but also lengthens the rotational correlation time and increases local contrast agent concentration, leading to improved image contrast. Using stimulus-responsive polymers as contrast agent carriers can activate the contrast agents selectively in pathological tissues, where the microenvironment conditions such as pH and ion concentration are different from those in the normal tissues. Therefore, the responsive polymer functionalized MRI contrast agents have improved relaxivity and retention time in tissues and offer a contrast between a diseased tissue and a normal tissue.15,106

pH-Responsive polymers have great potential in MRI applications since the pH in pathological regions (e.g. pH in tumor tissues is around 6.8) is lower than that in normal tissues (pH 7.4). In mildly acidic environments, the pH-sensitive polymers respond in different manners, such as a volume/conformation change or degradation.

Kikuchi and co-workers have designed Gd(III)-based MRI probes with Gd attached to a pH-sensitive n-octylamine-modified poly(sodium maleate-alt-ethyl vinyl ether) (poly(SM–EVE). When the surrounding pH decreased from 10 to 4, the conformation of the polymer changed from an extended state to a compact globular state due to the protonation of the amine groups. This conformation change was expected to induce rotational mobility of the Gd(III) complexes, which in turn influenced the relaxivities of the MRI probes. It was found that this decrease in pH increased the rotational correlation time from around 3 ns to 13 ns. Because the rotational correlation time is an important parameter affecting the relaxivities of the contrast agents, this type of MRI probe could be useful for the detection of diseases such as cancers.107

Linearly and spherically structured poly(methyl methacrylate) (PMAA)-based polymers were conjugated with Gd contrast agents to study their responsiveness to pH and their performance as contrast agents. The systems showed increased longitudinal and transverse relaxivities (r1 and r2, respectively) at a reduced pH. For instance, at pH 4, the polymers showed enhanced r1 values (28.0 and 12.1 mM−1 s−1 for the spherical and the linear polymers, respectively) compared to the r1 values at pH 7 (13.6 and 6.7 mM−1 s−1 for the spherical and the linear polymers, respectively). A spherical polymer system containing N,N′-methylenebisacrylamide crosslinker exhibited enhanced relaxivities compared to those of the linear PMAA polymers because of the reduced movement of the polymer molecules and the shorter distance between the adjacent Gd ions.108

Gd(III) complexes have been attached to poly(amino acid), where the relaxivities increased with increasing pH. Poly(amino acid) is flexible in an acidic medium, but the polymer flexibility is reduced with increased pH because the intrachain hydrogen bonds lead to a rigid conformation. Therefore, the polymer-coated Gd(III) contrast agents exhibit a change in relaxivity of the Gd complex. The relaxivity of the contrast agent increased from around 23 to 32 mM−1 s−1 when the pH increased from 4 to 8.109

Ultrasmall gadolinium oxide nanoparticles have been encapsulated by a biodegradable stimulus-responsive poly(lactic-co-glycolic acid) (PLGA) to increase the efficiency of MRI and enable the detection of multiple chemical species. A decrease in pH or increase in reactive oxygen species (ROS) concentration accelerated the degradation of the polymer matrix and released the contrast agent which can then be detected by MRI.110

A pH-responsive theranostic nanoprobe composed of Fe3O4 nanoparticles (T2 contrast agent), Mn-porphyrin (T1 contrast agent) and a layer of poly(acrylic acid) on silica (PAA-SiO2), loaded with contrast agents and a model antitumor drug, doxorubicin (DOX), was designed as an MRI contrast agent in cancer tissues. The Fe3O4 nanoparticles were coated with a layer of SiO2 that interacted with PAA through hydrogen bonds between the hydroxyl groups on silica and the carboxyl groups on PAA. The PAA was functionalized with a tumor targeting group, cRGD, while DOX and the Mn-porphyrin contrast agent were encapsulated inside the nanocomposite. In cancer tissues, the carboxyl groups of PAA were protonated, weakening their interaction with positively charged DOX. Lowering the pH from 7.4 to 5.0 caused the ratio r1/r2 to decrease, resulting in an increase in T1 contrast enhancement as the Mn-porphyrin was released. Hence, the pH-sensitive contrast reagents allowed the T1/T2-dual mode contrast agent to provide more accurate diagnostic information.111

Gao and co-workers have reported a new class of contrast agent based on a pH-responsive block copolymer, poly(ethylene glycol)-b-poly[2-(diisopropylamino)ethyl methacrylate] (PEG-b-PDPA), for MRI applications. PEG-b-PDPA formed micelles at pH 7.4 and dissociated into unimers at a pH lower than 6.3. This change arose from the protonation of the tertiary amine groups, causing the polymer to change from hydrophobic to hydrophilic. The transition pH of the ionizable polymer was close to the physiological pH, making this copolymer interesting for MRI applications via a chemical exchange saturation transfer (CEST) mechanism. The CEST mechanism is a contrast mechanism based on the proton exchange between the labile solute protons and the bulk water. When the exchangeable proton spins are saturated, the saturation is transferred to the bulk water via a chemical exchange, which enhances the detection sensitivity. The polymer-based CEST probes did not exhibit any signal at physiological pH, since the micelles did not have any exchangeable protons. However, in an acidic medium, the presence of the protonated unimer turned on the CEST switch, allowing the detection of acidosis-related metabolic diseases.112

1H, despite being the most widely used for cancer diagnosis, is associated with drawbacks such as insufficient sensitivity towards tumor tissues because of its presence in healthy tissues. Alternatively, 19F of 100% natural abundance has a sensitivity comparable to 1H and is not present in the human body. Hence, 19F MRI produces high contrast images with improved signal to noise ratios. It is required that drugs and contrast agents be modified to be detectable by 19F MRI. A pH-responsive PEGylated composite was used as a fluorine probe where the PEGylated nanogel consisted of a polyamine core that underwent a hydrophobic–hydrophilic transition with the environment pH change (Fig. 2a).76 This volume phase transition influenced the mobility of the fluorinated compounds in the core of the nanogel, and altered the intensity of the 19F NMR signal. The poly(2,2,2-trifluoroethyl methacrylate)-tethered PEG, PEGylated PTFEMA, showed no fluorine signal observed at pH 7.4 by 19F NMR (Fig. 2b and c), but generated clear F signals at pH 6.5. Under physiological conditions, the T2 value (<1 ms) was too short to generate any signals in NMR. However, the T2 value (57 ms) was significantly increased at pH 6.5. Therefore, the PEGylated nanogel could be used as a tumor-specific 19F nanoprobe because the T2 value was regulated by changing the pH of the medium.75,77

image file: d0tb00366b-f2.tif
Fig. 2 (a) Schematic illustration of the smart 19F nanoprobe based on the pH-sensitivity PEGylated nanogel that swells with reduced pH. (b) pH dependency of the intensity of the 19F MR signal from the PEGylated nanogel containing 10 mol% of TFEMA. (c) 19F NMR spectra of TFEMA at various pH values. Reproduced from ref. 76, with permission from the American Chemical Society, copyright 2007.

Besides pH, temperature-sensitive polymers have also found applications in MRI imaging. Magnetite nanoparticles (Fe3O4), a type of MRI class T2 contrast agent, were coated with a temperature-responsive copolymer containing PNIPAM and an acrylate monomer Nile red, a fluorescent dye, to produce a contrast agent.113 The copolymer, adsorbed onto the magnetite, caused changes in relaxation times, which were investigated via fluorescence measurement. Below LCST (at T = 25 °C), the polymer chains exhibited repulsion and expanded in the solution, making a stable nanoparticle dispersion. When the temperature was raised above LCST (at T ∼ 33.5 °C), the polymer chains became attracted to one another due to dehydration of the polymer, causing aggregation of the nanoparticles and an increase of hydrodynamic size from 38 nm (T = 25 °C) to more than 1 mm (T = 40 °C). Fig. 3a illustrates the temperature dependency of the aggregation of PNIPAM coated magnetite nanoparticles, with a more compact conformation above the LCST. The responsiveness of the nanoparticles to temperature and magnetic field is shown in Fig. 3b. At 25 °C (T < LCST), the particles existed in a highly dispersed state, observed as a homogeneous solution, even in the presence of a magnetic field. As the solution temperature was increased to 40 °C (T > LCST), the particles aggregated, causing a sharp increase in the turbidity of the solution. Followed by the aggregation of the particles, sedimentation took place in the presence of a magnet. In the case of PNIPAM-co-Nile red adsorbed on magnetite nanoparticles, the complex aggregated at 35 °C, evidenced by fluorescence quenching. Hence, the T2 relaxation time of the PNIPAM-coated NPs reached a peak around 0.05 s at 35 °C and this value decreased as the temperature of the milieu was increased. The thermo-responsive polymer associated with magnetic core provided a multi-stimulus-responsive nanoparticle.113

image file: d0tb00366b-f3.tif
Fig. 3 (a) Lower critical solution temperature (LCST)-induced aggregation in water of individually coated magnetite-PNIPAM nanoparticles. (b) Dispersion–flocculation behavior of magnetite-PNIPAM nanoparticles, as a function of temperature and magnetic field (concentration = 20 mg mL−1). Reproduced from ref. 113, with permission from the American Chemical Society, copyright 2011.

PNIPAM modified MRI contrast agents have been used to investigate the sensitivity of manganese tetra(3-vinylphenyl)porphyrin to temperature changes using the PNIPAM-N,N′-methylenebisacrylamide (PNIPAM-MBA) crosslinker. The system demonstrated a sharp hydrodynamic change upon varying the temperature around the LCST of the polymer composites (29–33 °C). For instance, for a microgel containing 0.17 wt% Mn, the hydrodynamic diameter decreased from 530 nm to 335 nm when the temperature was raised from 25 °C to 37 °C, with a sharp phase transition in the narrow temperature range of 29–35 °C. The shrinkage increased the density and the steric effect of the groups inside the microgel. As a result, the paramagnetic core motion was reduced. The change in the volume of the microgel affected the longitudinal relaxivity, r1 (for 0.17 wt% Mn microgel, r1 decreased from 14.4 to 12.3 mM−1 s−1 as T increased from 25 °C to 37 °C) and hence influenced the image signals. Hence a rise in the temperature resulted in the enhancement of MRI relaxivity and can hence be applied in imaging where a small change in the temperature can be monitored by MRI.114

Ion-responsive polymers have been investigated in cancer detection because the ionic composition of cancerous tissues differs from that of normal tissues. An ion-responsive copolymer (poly(OEGMA-co-TFEA)) with labeled fluorine was studied for a potential application as a MRI contrast agent.115 A simulation of the copolymer in pure water and in a salt solution illustrated that the hydrodynamic diameter of the copolymer decreased and the polymer adopted a more globular conformation in a NaCl solution because the ions disrupted the interactions between water molecules and OEMGA (Fig. 4a and b). Fig. 4c shows the simulation of the gyrational radius (Rg) of the copolymer in water and a 12.4 wt% NaCl solution in 30 nanoseconds, showing that the Rg of the copolymer is larger in water than that in a salt solution. The polymer conformation changes with various salt concentrations (from 0 to 0.5 M NaCl) resulting in the relaxation time T2 change in the 19F NMR of the copolymer (from 180 to 130 ms). Since the relaxation time T2 is determined by the 19F nuclei dipolar coupling strength with fluorine and proton nuclei in the proximity, the presence of salt affected relaxation times T2 by changing the polymer conformation (Fig. 4d). The 19F NMR T2 of the copolymer of the normal breast cells (MCF-12A) and breast cancer cells (MCF-7) showed that the polymer had short T2 relaxation times in the cancer cells with higher ionic content.115

image file: d0tb00366b-f4.tif
Fig. 4 Snapshots of the poly(OEGMA-co-TFEA) chains taken at the end of ∼30 ns simulation in (a) pure water (backbone carbon atoms are shown in blue) and (b) salt solution (12.4 wt% NaCl) (backbone carbon atoms shown in red). Water molecules and NaCl ions have been omitted for clarity. (c) Radius of gyration (Rg) of the poly(OEGMA-co-TFEA) copolymer in pure water in the presence of 12.4 wt% NaCl during ∼30 ns simulation using LAMMPS at 298 K. (d) 19F NMR T2 relaxation times for the poly(OEGMA-co-TFEA) copolymer in deionized water and in NaCl solutions (0.1 to 0.5 M) at 310 K measured at a field strength of 9.4 T. Reproduced from ref. 115, with permission from the American Chemical Society, copyright 2016.

Light was reported as a stimulus in the study of the photosensitive MRI contrast agent based on superparamagnetic iron oxide nanoparticles (SPIONS) and amphiphilic poly(ethylene oxide)-b-poly(2-((((2-nitrobenzyl)oxy)carbonyl)amino)ethyl methacrylate) (PEO-b-PNBOC) copolymer. The copolymer could self-assemble into vesicles and a UV irradiation transformed NBOC into 2-aminoethyl methacrylate (AEMA), causing a transition from hydrophobicity to hydrophilicity.103 The hydrophilic vesicles allowed the release of the encapsulated contrast agent for their detection by MRI. This new photoresponsive polymer-containing a T2-contrast agent provided enhanced MRI constrast.93

Dual-responsive systems were also reported for theranostics that involved simultaneous MRI diagnosis and therapy. A theranostics design for imaging of cancer cells and drug delivery was produced using a dual temperature- and pH-responsive polymer, magnetic nanoparticles tethered with targeting agent, folic acid (FA) and a fluorescence imaging dye (rhodamine isothiocyanate).116 Polyethylene imine-based dual-responsive copolymers were covalently tethered to citrate stabilized monodispersed mixed ferrite nanoparticles (MFNPs) (CA-MFNPs). Folic acid was attached to the polymers to grant the cancer targeting property. DOX was loaded onto the nanoparticles by suspending them in a DOX solution, when DOX was entrapped in the polymer through hydrophobic interactions and hydrogen bond formation. It was demonstrated that nanoparticles with FA induced higher cell death of HeLa cells expressing folate receptor than those without FA. The performance of the nanoparticles as MRI contrast agents was evaluated through in vitro MR imaging of folate receptor positive HeLa cells. The T2-weighted phantom images exhibited a significant negative contrast enhancement, indicating the efficacy of the nanoparticles as MRI contrast agents.116

A theranostic composite nanoplatform with a multimode therapeutic effect (photothermal/photodynamic therapy/chemotherapy) and dual-mode imaging function (MRI/X-ray CT) was produced based on a mesoporous silica drug delivery system.14 Different functional groups were added to the mesoporous silica nanoparticles to obtain targeted functionality. The spheres were modified with amino groups and then chelated with Gd (MRI contrast agents) followed by the coupling of Chlorin e6 (Ce6, a photosensitizer for photodynamic therapy) using the condensation reaction of the amide bonds. The chemotherapy drug DOX was encapsulated into the channels of the porous spheres of the silica. The thermo/pH-sensitive polymer poly[(N-isopropylacrylamide)-co-(methacrylic acid)] (P(NIPAM-co-MAA)) was encapsulated in the pores of the silica spheres, which acted as a “gatekeeper” to control the release of the DOX due to temperature change. In addition, carbon dots (CDs) were attached to the platform. These carbon nanoparticles could absorb the photon energy from a 980 nm laser and then convert the irradiation to heat. The generated heat increased the local temperature to shrink the temperature-responsive P(NIPAM-co-MAA), opening the “gate” to release the DOX.14 The in vivo and in vitro studies have demonstrated that the system provided enhanced MRI signals and more efficient chemotherapy.

3.2. Capture and purification of biomolecules through protein–ligand recognition

Molecular recognition is a fundamental process in living organisms that involves highly specific interactions between biomolecules to generate specific complexes. The interactions between proteins with ligands are crucial for drug development and medical diagnosis. The conjugation of proteins and ligands with stimulus-responsive polymers has been reported to facilitate processes such as affinity separation, through reversible phase changes associated with the responsive polymers.117,118

Streptavidin, a tetrameric protein that can bind to biotin molecules with a high affinity, has been widely used as a biomarker protein. Conjugation of streptavidin with PNIPAM was reported by Stayton and co-workers as a system to control the ligand binding affinity of proteins.118 The temperature-sensitive polymer could manipulate the accessibility of biotin to the binding sites on the bioconjugate. Heating the bioconjugate caused the polymer to collapse into a globular conformation that blocked the access to the binding site. Upon cooling the system, the hydration of the polymer opened the access to the binding site, allowing the biotin to bind to streptavidin. Attaching a genetically engineered streptavidin to PNIPANM allowed for a specific reversible binding and release of biotin controlled by temperature.119

The conjugation of oligonucleotides to PNIPAM and the protein allowed for a selective and sequential target molecule isolation from complex mixtures under different conditions. The high affinity of streptavidin to biotin allowed the streptavidin–PNIPAM complex to bind to radiolabeled biotin/biotinylated alkaline phosphatase. The PNIPAM moiety underwent phase separation when the solution temperature was increased to 32 °C, resulting in the precipitation of the biotin–streptavidin–PNIPAM complex.120

Nash and co-workers reported an approach to concentrate streptavidin in spiked human plasma using gold nanoparticles (AuNPs) modified by a diblock copolymer containing a thermally responsive PNIPAM block, a cationic amine-containing block, and a semi-telechelic PEG2-biotin end group.121 As depicted in Fig. 5, the PNIPAM-AuNPs were mixed with PNIPAM-Fe3O4 nanoparticles (mNPs) at room temperature (<LCST of PNIPAM) and the mixture was then brought in contact with streptavidin spiked human plasma. The streptavidin bonded to the biotin end group on PNIPAM-AuNPs. The solution was then heated to a temperature higher than the LCST of PNIPAM when PNIPAM-AuNPs and mNPs generated magnetic aggregates because of the collapse of PNIPAM. The magnetic aggregates were extracted from the human plasma using an external magnet. The aggregates were concentrated into a 50-fold smaller fluid volume at room temperature when the gold nanoparticles bound with the streptavidin target were dispersed at a temperature below the LCST of PNIPAM. The concentrated gold-labeled streptavidin was subsequently analyzed directly using lateral flow immunochromatography.121

image file: d0tb00366b-f5.tif
Fig. 5 Schematic illustration of concentrating streptavidin using gold nanoparticles and magnetic nanoparticles with PNIPAM. AuNPs coated with biotinylated diblock copolymers bind to streptavidin spiked into 50% human plasma. mNPs coated with homo-PNIPAM are added, and the temperature is raised above the polymer LCST. Mixed streptavidin-AuNP/mNP aggregates are separated by a magnet. The separated aggregates with gold-labeled target protein are resuspended into a smaller volume of a buffer below the LCST and flowed through an LFIA strip with immobilized capture antibodies. Visualization of the target protein is achieved by AuNP light extinction at the capture line of the LFIA strip, while the non-biofunctional mNPs are rinsed away. Reproduced from ref. 121, with permission from the American Chemical Society, copyright 2010.

Poly(NIPAM-co-AAc), a thermo- and pH-responsive copolymer, was conjugated with genetically engineered streptavidin to form a bioconjugate complex for diagnostics. The combination of temperature and pH responsiveness generated a synergistic effect due to the hydrogen bonding between the AAc and NIPAM units. At pH 7.4, the LCST of the copolymer (52 °C with 6 mol% AAc in copolymer) was higher than that of PNIPAM homopolymer (32 °C) because of the presence of COO.84,85 It was found that the copolymer was in a collapsed conformation at low pH (4.0) and high temperature (37 °C). The copolymer expanded when the pH increased to 7.4 due to the repulsion between COO groups. When this copolymer was conjugated to a specific site near the biotin-binding site of streptavidin, biotin binding at 37 °C was significantly reduced at pH 4.0 compared to that at pH 7.4 because the collapsed copolymer at pH 4.0 blocked the biotin binding site. This design allowed the reversible dual-responsive binding of biotin to streptavidin, which could be applied for medical diagnostics in body parts where the pH is quite low, such as in salivary glands and the stomach.85

Immunoglobin, IgG, is an important antibody used to detect diseases like autoimmune hepatitis. Several systems have been developed to separate the antibody through antibody–antigen interactions. PNIPAM has been used to facilitate the separation of IgG via different methods. Hoffmann and co-workers conjugated PNIPAM with IgG to interact with specific antigens to form an antibody–antigen–PNIPAM complex, which precipitated out of the solution at high temperature (T > LCST). This phase separation allowed the determination of bound antigen that can be useful in enzyme immunoassays for IgG.122

PNIPAM was also conjugated with protein A, a major cell wall component of the Gram-positive bacterium Staphylococcus aureus. The polymer composite interacted with IgG to form a complex that precipitated when the solution was heated to 37 °C. This affinity precipitation was found to be a good method for the separation of IgG and can hence be applied for medical diagnostics.123

Combining IgG with PNIPAM using 3-mercaptoporpionic acid has produced IgG-PNIPAM bioconjugates that responded rapidly to solution temperature changes. After binding to an antigen, the antigen–IgG-PNIPAM complex also exhibited temperature-sensitivity. A rapid phase separation took place when the temperature of the surroundings rose above LCST of the conjugate (∼33–34 °C), allowing the separation of IgG.124

Oligo-NIPAM (ONIPAM), which is highly soluble in water at a low temperature, exhibits a phase separation near 32 °C (LCST). Grafting ONIPAM to atelo collagen, a component of diagnostic assays, generated temperature-responsive ONIPAM–collagen bioconjugates. The ONIPAM–collagen precipitated from its aqueous solution when the solution was heated to 34 °C, allowing the isolation of the bioconjugate by a small temperature change.8

Human genomic DNA (gDNA) is a disease index for breast cancer and has the potential for rapidly diagnosing breast cancer recurrence after surgery. A method to capture gDNA from human blood incorporating target DNA has been designed using pH-responsive poly(2-dimethylaminoethyl methacrylate) (PDMAEMA) tethered onto a silicon surface. At pH 6.0, the DNA was adsorbed onto the hydrophilic PDMAEMA via hydrogen bonding between the positively charged PDMAEMA and the DNA. The release of gDNA from the surface was achieved by increasing the pH to 8.0, reducing the positive charges on the polymer, leaving a more hydrophobic surface. The released gDNA was amplified through a polymerase chain reaction and analyzed through agarose gel electrophoresis.83

Immunoaffinity purification of antibodies was achieved by using magnetic poly(styrene/N-isopropylacrylamide/methacrylic acid) latex particles, which exhibited a reversible transition between dispersion and flocculation upon changing the temperature and the concentration of NaCl in solutions. The temperature-sensitivity of the particle was attributed to the NIPAM moiety. It was observed that increasing the temperature led to a decrease in the critical flocculation concentration, a result of the dehydration of the NIPAM layers of the surface. The interesting thermal and magnetic properties and large surface area of the particles made them useful for affinity purification. These thermo-sensitive magnetic latex particles were used to covalently immobilize antigen bovine serum albumin (BSA), a disease marker for hepatitis E virus infection, to produce immunomicrogels. The microgels were used to bind anti-BSA antibodies in antiserum through the interactions between BSA and anti-BSA antibodies. The obtained complex microgels were precipitated at high temperature and separated by a magnetic field. This study provided a method to separate and purify anti-BSA antibodies from antiserum, by controlling the temperature and magnetic field.114,125

4. Polymer lab-on-a-chip for medical diagnostics-mechanisms and applications

The lab-on-a-chip devices belong to the family of microelectromechanical systems, a technology to miniaturize processes on an integrated system.126–128 (Bio)chemical processes that take place on lab-on-a-chip devices usually use the microfluidic technology that utilizes microchannels allowing manipulation of very small volumes of up to a few picoliters. Lab-on-a-chip, when used in medical diagnostics, offers many advantages such as reduction in cost, time, reagents and samples required, as well as better sensitivity and parallelization resulting from device miniaturization. On the other hand, small size has imposed many challenges in the device fabrication. For example, macroscale fluid control parts like pumps, valves and mixing components need to be reduced to microns to be integrated in lab-on-a chip devices. It is obvious that fluid control in lab-on-a chip devices cannot rely on mechanical components but utilize responsive properties of building blocks. Stimulus-responsive polymers have demonstrated great potential in building fluid control components129,130 and capturing diagnostic targets through specific interactions with biomarkers.20,32–34,36,131

4.1. Temperature and light controlled actuators

Temperature and light are the most common stimuli for passive actuators in lab-on-a chip devices because they can provide localized changes in responsive polymers. Temperature-responsive hydrogels are often used as valves that open and close by temperature changes in microfluidics either through the change of hydrophilicity/hydrophobicity or the swelling/deswelling of hydrogels. A fully integrated microfluidic valve with a switchable, thermo-sensitive polymer surface has been fabricated through the combination of polyelectrolyte multilayer films, thermo-responsive polymer brushes and microfabrication. To build a thermosensitive valve, a patch of microfluidics was conformally coated with polyelectrolyte and silica nanoparticle multilayer coatings followed by the grafting of PNIPAM from the polyacrylic acid on the top of deposited films. When the valve was heated above LCST, the valve (PNIPAM) became hydrophobic and inhibited the flow of water. In contrast, as the valve was cooled down to room temperature, it became hydrophilic and allowed the flow of water.55

When temperature-responsive hydrogels were used as valves, the collapsed hydrogel blocked part of the microchannel when the temperature was above LCST, granting an open state of the channel. Lowering the temperature below the LCST led to a swelling of the hydrogel and closing of the channels.18,132

Adding acrylamide (AAm) to the PNIPAM network provided P(NIPAM-co-AAm) hydrogels that swelled/deswelled over a broader temperature range (from 28 to 60 °C). The acrylamide in the block copolymer increased the concentration of hydrogen bonds with water molecules. Unlike PNIPAM that has a sharp LCST, the phase separation of the block copolymer happened over a broader temperature range (28–60 °C). Because more thermal energy is needed to break the hydrogen bonds, the total collapse of the network took place at 60 °C.133

Tabeling and co-workers have designed microfluidic devices with temperature controllable actuators:microcages for the retention/release of cells and microvalves for the opening/closing of the channels.18 These structures were fabricated by photopatterning PNIPAM onto the bottom of the microfluidic device. Once the hydrogel was built onto the surface, the device was closed using polydimethylsiloxane (PDMS) and NOA (Norland Optical Adhesive). A flow-rate monitor was introduced into the device to determine the hydrodynamic resistance of the device at various temperatures. The square microcages had a 50 × 50 μm base and 20 μm high wall that could trap a single cell. This allowed single cell manipulation, which increased the sensitivity of the sample analysis. The chamber was injected with a buffer containing the Human Synaptonjanin 1 (SYNJ1) gene and an amplification mix consisting of primers targeting the gene, at 40 °C (T > LCST). As the temperature was lowered to 25 °C, the hydrogel walls of the cage swelled up to trap the cells. This process was reversible. When the temperature was increased above LCST, the hydrogel walls shrank and released the cells. Two types of microvalves were built to control the flow in this study. The type one valve opened or closed the channel by heating/cooling (Fig. 6a). The type two microvalve provided more control over the operation of the device with local temperature at each valve, allowing the independent opening/closing of each valve (Fig. 6b). The method allowed two different fluids to be transported to the main channel with control over the ratio of the two fluids in the channel and time of release. The type 1 valve (Method 1) was tested for its reproducibility and negligible leaks were found after subjecting the valve to high pressure (500 mbar). The system described above was used to perform isothermal amplification and NAAT (nuclear acid amplification test). The fluorescence intensity of the cages was monitored using fluorescence probes. Comparing the intensity of fluorescence signals obtained from the system with the control experiment (DNA-free sample) allowed successful detection of the gene. The system was able to detect the SYNJ1 gene, whose mutation indicated several neurodegenerative diseases like Parkinson's disease. Variation of the different functions of the device showed the potential for expansion for the detection of other diseases.

image file: d0tb00366b-f6.tif
Fig. 6 (a) Temporal evolution of the flow-rate Q in a 9.4 μm × 400 μm × 1.2 cm channel integrating a hydrogel valve, using Method 1, and subjected to a temperature field switching between 25 and 50 °C, at a frequency of 1.4 mHz. Q0 is 7 μL min−1. (b) Device with two locally actuated hydrogel valves, elaborated according to Method 2. The channel widths are 400 μm (i): with the upper valve open and the lower valve closed, the red dye flows downstream. (ii) With the upper valve closed and the lower valve open, the green dye passes through. (iii) Typical fluorescence image for two valves alternatively open and closed at a 2 Hz frequency (Method 2). (iv) Evolution of the interface position Ai, that is, its distance from the symmetry axis of the collecting channel (see iv, inset), and divided by half the channel width w ≈ 400 μm. Ai is measured as a function of time, with the lower (green) valve staying open, and the upper (red) valve, subjected to a sequence of heating events. The (red) valve is initially heated above 32 °C. Then at t = 4.2 ± 0.1 s, heating is turned off; finally, at t = 8.9 ± 0.1 s, heating is turned on again. From (iv), one estimates that the closing and opening times of the valve are, respectively, 0.6 ± 0.1 s and 0.25 ± 0.15 s. Reproduced from ref. 18, with permission from Nature.

Beta and co-workers have used paper-based microfluidic technology for the detection of Escherichia coli (E. coli).133 The system used P(NIPAM-co-AAm) hydrogels as reservoirs to release fluids that would trigger the reactions between enzymes and antibodies. In this study, horseradish peroxidase (HRP) was used as the enzyme to react with E. coli antibody. The HRP was stored in the paper channel under dry conditions and upon trigger (T > LCST, which is 31 °C), the liquid in the hydrogel was released and flew onto the paper substrate due to capillary force. A hydrogel-driven assay was designed to detect E. coli infections in tap water based on a polyclonal antibody against the K99 pili of E. coli. The antibody was labelled at the NH-terminus with the HRP enzyme. Fig. 7a displays the layout of a paper fluidic device with the description of the stepwise operation, from the collapse of the hydrogel to the release of the E. coli antibody to the visual detection. The paper channels were preloaded with antibody, barium peroxide (BPO) and tetramethylbenzidine (TMB). TMB is a substrate for horseradish peroxidase, producing a soluble blue reaction product. A water loaded hydrogel pad was placed on the reservoir field of the inner paper channel and a second hydrogel pad loaded with 100 μM sulfuric acid was placed on the reservoir field of the outer circular paper channel. The micro-pore filter was exposed to the sample liquid and dried at room temperature. To test the function of the assay, 1 mL of an E. coli containing water test sample with a cell density of 2.8 × 10−5 cells per mL was passed through the micro-pore filter. Afterwards, the filter was loaded into the test assay between the channel layer and the reservoir paper. To test the sample, the water containing hydrogel was first heated up to 38 °C to release half of its water content and the heating was stopped. The released water was drawn into the paper channel by capillary force. The water dissolved the E. coli antibody in the channel, and the fresh antibody solution reached the micro-pore filter (sample paper) as shown in Fig. 7a(1). The same hydrogel was then heated up to 38 °C again to release the remaining water, which washed the channel structure and the micro-pore filter as shown in Fig. 7a(2). If the target bacteria were present on the micro-pore filter, the labelled antibody attached to the pili of the microorganisms. The second hydrogel pad containing sulfuric acid was heated up to 38 °C, so that the BPO and the TMB were dissolved and also directed to the micro-pore filter via the outer ring-shaped paper channel. If microorganisms marked by the antibody were present, the color changed from yellow to blue as shown in Fig. 7b. In the absence of E. coli contamination of the micro-pore filter (control experiment), the antibody was washed to the paper pad in the bottom layer below the micro-pore filter and no color reaction was initiated by the release of BPO and TMB. Fig. 7b and c are the photographs of the E. coli test chips for a positive and a negative response, respectively.133

image file: d0tb00366b-f7.tif
Fig. 7 Operation of the hydrogel-driven paper fluidic E. coli test assay. (a) Flow chart illustrating the steps of operation of the device, (1) release of the antibody, (2) washing of the sample, and (3) release of the color indicator. Left and right hand sides show positive and negative tests respectively. (b) Photograph of a paper-based microfluidic E. coli test chip showing a positive response, indicated by the colored ring. (c) Photograph of a paper-based microfluidic E. coli test chip showing a negative response. Panels (a–c) were adapted and reproduced from ref. 133 with permission from the Royal Society of Chemistry.

Photoresponsive passive micromixers were produced from the light responsive hydrogel discussed previously.99 As shown in Fig. 8a, the light triggered a ring closing reaction to convert hydrophilic merocyanine to a hydrophobic apolar spiropyran. The release of the bonded water caused the hydrogel to shrink. The hydrogel was patterned on a microfluidic channel to create grooves for fluid mixing (Fig. 8c and d). In the dark, the gel was hydrated and grooves were exposed to fix the fluids from two inlets. When the hydrogel was exposed to blue or white light, the ring closing reaction caused the hydrogel to collapse, losing the groove features to mix fluids.

image file: d0tb00366b-f8.tif
Fig. 8 (a) Chemical structure of the monomers used to fabricate the light responsive hydrogel. (b) Photoisomerization of protonated merocyanine to spiropyran in an acidic environment and the corresponding thermal back reaction. (c) Schematic representation of photoresponsive passive micromixers based on a light responsive hydrogel. In the off-state the hydrogel is swollen, resulting in a passive mixer, while upon exposure to light, the gel shrinks resulting in a nonmixing flat state. (d) Schematic overview of the micromixer. Reproduced from ref. 99. Copyright 2017, John Wiley and Sons.

4.2. Capturing diagnostic targets

Lab-on-a-chip technology has facilitated the isolation and detection of diagnostic targets from complex solutions by using responsive polymers. By combining antibody–antigen affinity with stimulus-responsive polymers on a single chip, new diagnostic systems have been developed.

Diagnosis of malaria is usually carried out via the detection of the antigen that is released by the red blood cells due to the presence of the parasite in the blood. While the available tests allow the detection of high levels of antigen in the blood, they are not suitable for lower concentrations of antigens.134 A temperature controlled microfluidic card was developed to control the passage of particles, and allowed the purification, concentration and detection of the targeted antigen.135 The card contained a porous PNIPAM-grafted nylon-6,6 membrane positioned in multilayer microchannels composed of poly(ethylene terephthalate) (PET) and poly(methylmethacrylate) (PMMA) (Fig. 9). PNIPAM was modified with tetrafluorophenol to yield amine-reactive ester groups for conjugation to amine groups of anti-streptavidin and anti-PfHRP2 (Plasmodium falciparum histidine-rich protein 2) antibodies. The modified PNIPAM was mixed with sample solutions to bind antigens to PNIPAM. The liquid was pushed through the porous PNIPAM-grafted nylon-6,6 membrane using a syringe pump. Heating the card to 35 °C (T > LCST) allowed the liquid to flow through the membrane, while the antibody conjugate complex with modified PNIPAM underwent phase separation at the membrane. Upon cooling the card to 25 °C (T < LCST), the PNIPAM chains were solvated and the antibody complex redissolved in the flow and passed through the membrane. All flow fractions exiting the channel were analyzed by a plate reader fluorimeter, revealing that the antibody–PNIPAM conjugate could be purified from a complex solution, by a simple manipulation of the temperature. This concentrator system was successfully applied for the detection of the malaria antigen PfHRP2 from spiked human plasma and the results were comparable with traditional ELISA-based tests with a much faster detection time.135

image file: d0tb00366b-f9.tif
Fig. 9 Schematic system design for immunocomplex capture and concentration under conditions of capture (>LCST) and release (<LCST). Covalently modified PNIPAM–antibody, bound antigen, and detection antibody are captured at the membrane during polymer aggregation above the LCST, while unconjugated plasma components flow through. When the membrane region of the device is cooled, the polymer becomes hydrophilic and molecular conjugates are released back into the flow stream, carrying their cargo of antigen and detection antibody with them. Reproduced from ref. 135, with permission from the American Chemical Society, copyright 2010.

A lab-on-a-chip device was designed to capture diagnostic targets using a dual-stimulus-responsive nanoparticle system as shown in Fig. 10a.136 Magnetic nanoparticles (mNPs), γ-Fe2O3, were used as the core and PNIPAM chains, modified with dodecyl chains and carboxylate moieties self-assembled around the core to form PNIPAM mNPs. Biotin was attached to the PNIPAM mNP complex through the carboxyl group of the polymer, generating b-PNIPAM-mNPs that were introduced into the microchannel. In the PEGylated microfluidic channel, b-PNIPAM-mNPs exhibited a low magnetophoretic mobility when the temperature was below LCST (LCST of PNIPAM-mNPs = 32.4 °C). The b-PNIPAM-mNPs flew freely in the channel and readily bound to the streptavidin biomolecules. When the complex streptavidin–biotin-PNIPAM mNPs flew into the heated section of the microchannel where the temperature was above LCST (LCST of streptavidin–biotin-PNIPAM-mNPs = 41.1 °C), they formed aggregates. When a magnetic field was applied, it caused these aggregates to immobilize on the microchannel wall. The immobilization of the conjugate complex facilitated the washing process since any mobile species was flushed out of the channel. Removal of the magnetic field and cooling the channel below LCST resulted in the redissolution of the aggregates and their elution through the channel (Fig. 10a). The capture/release of PNIPAM-mNPs is observed in the micrographs in Fig. 10b. This dual stimulus-responsive system allows the isolation of diagnostic targets that can be used in point-of-care devices.136

image file: d0tb00366b-f10.tif
Fig. 10 Particle capture and release scheme (a) and the corresponding micrographs (b). The PNIPAM-mNP capture/release was demonstrated in PEGylated PDMS microfluidic channels whose channel width was 500 μm. The magnetic field was introduced by embedding a magnet at the lower side of the channel. The mNP solution (4 mg mL−1) was injected into the channels with a constant flow (∼1 μL min−1) during the entire experiment. The mNPs are soluble and freely flowing in the PEGylated channels when the temperature is below the LCST of PNIPAM. As they flow into the heated region, the temperature is above the LCST of the PNIPAM and the mNPs aggregate, but do not stick to the non-fouling PEGylated channel walls in the absence of an applied magnetic field. The mNPs are captured onto the PEGylated channel walls only when the temperature is raised above the LCST and the magnetic field is applied. The reversal of the temperature and applied magnetic field results in the redissolution of the aggregated mNPs and their diffusive reentry into the flowstream. Reproduced from ref. 136, with permission from the American Chemical Society, copyright 2007.

Stayton and co-workers reported the application of a temperature-responsive polymer for protein–ligand recognition. Primary amine-functionalized polystyrene latex beads were covalently modified with PNIPAM and PEG-biotin, which introduced the biotin moieties. The isolation/purification process followed a similar pathway as described above. When a suspension of these “smart” beads was introduced into the microchannel of the chip, the beads flew through the channel at room temperature. The polymer underwent a hydrophilic–hydrophobic phase transition when the temperature of the device was raised above the LCST of the PNIPAM-coated beads (37 °C). As a result, the beads aggregated and adhered to the PET channel walls. The aggregates were purified by passing a buffer solution through the channel to remove the free beads. Following the purification step, a sample of streptavidin was introduced into the device. The streptavidin bound to the adhered biotin, generating streptavidin–biotin bound beads that remained adhered to the walls until the temperature was lowered to room temperature. The low temperature caused the release of the complex beads from the wall and their elution through the microchannel. Hence, the system was used to reversibly bind and release affinity molecules, allowing the detection of target molecules by simply changing the temperature of the device.137

The system established earlier by Lai136 was studied further for its application as an immunoassay system for quantitative analysis, using a digoxin antibody–antigen couple. PNIPAM and PEG modified latex beads were functionalized with streptavidin. Addition of biotinylated anti-digoxin IgG to the system resulted in the streptavidin–biotin conjugation on the antibody bound beads. These antibody-beads adhered to the PET walls of the microchannel at 37 °C due to the collapse of PNIPAM. A sample of the target antigen molecule, digoxin, and fluorescently labeled digoxigenin was introduced into the microchannel flow. Competitive binding of the antibody to digoxin and digoxigenin took place and upon cooling the system to room temperature, the complexes were dissolved and flew through the channel. The sample exiting the channel was analyzed to obtain the measurement of the target antigen bound to the antibody. This device provides a precise and reversible temperature-responsive system for the immobilization of the biomolecules and their analysis.138

Besides temperature, pH change has also been used to separate and purify biomarkers. A dual-responsive system was designed for biotin–streptavidin affinity separation.139 The pH-responsive mNP used in the system was prepared as follows: the copolymer P(tBMA-NIPAM) was first synthesized and then reacted with Fe(CO)5 with heating. The heating caused partial cleavage of the t-butyl group of the PtBMA to generate P(MAA-NIPMA) coated γ-Fe2O3 mNP micelles. The micelles were functionalized with biotin and streptavidin was added to the mixture, forming biotin–streptavidin pH-responsive mNPs. When the pH of the solution was brought to 7.3, the hydrophobic PMAA adopted a hypercoiled conformation, leading to aggregation of the conjugates. The solution was then introduced into the first microfluidic channel flowstream at the same pH of 7.3 (Fig. 11). With an external magnet, a magnetic field was applied, causing the aggregates to move under magnetophoresis laterally through the interface into a second flowstream. The pH of the second stream was higher (8.4), resulting in the redissolution of the aggregates, under continuous flow. The magnetic field did not affect the flow of the aggregates because the particle size changed when the nanoparticles redissolved, experiencing low magnetophoretic velocity.139

image file: d0tb00366b-f11.tif
Fig. 11 Target analyte separation in a microfluidic channel facilitated by pH-responsive mNPs under isothermal conditions. The channel contains two flow streams. The left stream (green) is the sample that has been pre-incubated with mNPs. mNP aggregation is induced by using a lower pH buffer in this sample flowstream. The pH of the right stream (pink) is chosen to reverse mNP aggregation. A rare-earth magnet provides sufficient magnetic field to attract the aggregates laterally into the higher pH flowstream. The conjugate aggregates move out of the sample flowstream and in to the higher pH stream, where they return to a dispersed state, carrying the bound target analyte with them. Movement of the other molecules across this interface is limited by diffusion due to the low Reynolds number (laminar) fluid flow. The scheme was adapted and reproduced from ref. 139 with permission from the Royal Society of Chemistry.

Hoffmann and co-workers modified the walls of microfluidic channels to make them more applicable for protein–ligand affinity-based separation and for diagnostic assays. The walls of PDMS-based microfluidic channels were hence modified by UV radiation-grafted polymerization of stimulus-responsive polymers. Temperature-responsive PNIPAM, temperature- and pH-responsive P(NIPAM-co-AAc) and non-fouling hydrogel polyethylene glycol diacrylate (PEGDA) were used to prepare the “smart” microchannels that possessed reversible hydrophilic–hydrophobic walls. The modified surfaces provided large surface wettability changes when the temperature and/or pH was changed. For instance, changing the pH of the channel from 4 to 6 caused a change in water contact angle on the surface from around 65° to 40°. Increasing the temperature of the microchannel from 20 °C to 40 °C led to an increase in water contact angle on the surface from less than 40° to more than 80°. These properties make the “smart” microfluidic channels suitable as valves and for reversible immobilization of target biomolecules for diagnostic applications.140

5. Conclusions & perspectives

Stimulus-responsive polymers contribute to medical diagnostics by changing the polymer structures and properties upon temperature and pH change, light irradiation, and specific interactions with biomolecules of interest. The understanding of the structure-property relationship of stimulus-responsive polymers has laid a solid foundation for applying these polymers to make diagnosis faster, simpler and more cost effective. Although much effort has been devoted, the use of stimulus-responsive polymers in medical diagnostics is still in a developing stage, and several research areas require significant efforts to advance the field. It is obvious that the choice of polymers applicable to medical diagnosis is very limited. Temperature-responsive PNIPAM and its derivatives are the major players. Polyelectrolytes are another choice for pH- and ion-responsive polymers. Other stimuli besides temperature, pH, ion and light need to be considered. The local level of ROS can be used to distinguish inflamed tissues or tumors from normal tissues because ROS levels in pathological tissues can be 2 to 3 orders of magnitude higher than that in normal tissues.141 Combining polymers containing ROS-sensitive groups such as thioketal, vinylditioether, etc. with imaging reagents will allow real-time and noninvasive diagnosis by responding to ROS concentration changes in vivo.43 Attaching active biomolecules to responsive polymers provides a versatile method to enhance the molecular recognition capability.

The sensitivity to environment change is crucial in medical diagnostics because the change of temperature and pH is small in vivo and in vitro. While synthetic polymers are responsive to certain types of environmental changes such as temperature, pH, ion concentration, and light, nanomaterials have been used to extend the scope of stimuli. For example, carbon nanoparticles have been used to convert light irradiation to heat that can induce the structural change of temperature-responsive polymers.14 Magnetic nanoparticles are combined with PNIPAM to make polymers responsive to temperature and magnetic fields.139 Combining stimulus-responsive polymers and nanomaterials will greatly improve their versatility and efficiency. However, their toxicity and the impact of polymeric and nanomaterials on human immunology require extensive investigation for in vivo applications like MRI. The response kinetics of stimulus-responsive polymers in vivo and in vitro has not been systematically investigated. How fast these polymers can change their structures is important for time-sensitive diagnostics.

The application of stimulus-responsive polymers in lab-on-a-chip technologies holds great promise for the future of medicine. Point-of-care diagnostics is an area that requires immediate attention and stimulus-responsive materials can contribute enormously in making efficient devices for rapid diagnosis. A recent review article has discussed the application of microneedles (1 to 100 μm in diameter) for transdermal diagnostics.142 These small needles can access biofluids beneath the skin in a nearly pain-free manner. While polymer hydrogels have been used as microneedles to collect fluid, responsive polymer hydrogels hold great potential as microneedle materials for diagnosis by converting polymer structural changes such as degree of ionization to electrical signals. It is clear that the collaboration among chemists, materials scientists, electrical engineers, and medical researchers will significantly advance the development of responsive polymer systems in medical diagnostics.

Conflicts of interest

There are no conflicts to declare.


D. A. thanks the Basic Science Research Program through the National Research Foundation of Korea (NRF) funded by the Ministry of Education (2018R1A6A1A03023788) for the financial support.

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