Keywords

1 Introduction

Controlled drug delivery systems have represented a huge improvement in the way of drug administration compared to conventional formulations (Noorian et al. 2020; Krause et al. 2019; Kim et al. 2019). These systems have the advantage that they can maintain the concentration of the drug in the blood and tissues in the therapeutic field, reduce the frequency of drug administration and increase patient comfort and compliance. However, there are several medical disorders (diabet, heart disorders, etc.) when the continuous release of drugs in shorter or longer periods is not appropriate. For these cases, the pharmaceutical formulations must release the drug when one of the parameters that regulates the body's functions changes. These systems are generally based on intelligent or stimuli sensitive polymers (Liu et al. 2019; Jamwal et al. 2019; Alsuraifi et al. 2018).

Smart polymers are a category of macromolecules that in aqueous solution suffer a phase transition when small changes of the external parameters occur. These parameters can be temperature (Lee and Bae 2020), pH (Sapre et al. 2020), ionic strength (Zhang et al. 2005), electric field (Shang et al. 2007), magnetic field (Manouras and Vamvakaki 2017), light (Sedlacek et al. 2019), redox potential (Singh et al. 2020), the presence of biomolecules (Fundueanu et al. 2013), etc. Among polymers sensitive to external stimuli, those sensitive to temperature and pH are the most used for biomedical applications because both parameters vary most frequently in the human body. Changes in temperature and pH cause modifications in the physico-chemical properties of polymeric pharmaceutical formulation that result in modulation of the release profiles of loaded drugs (Wang et al. 2020; Zheng et al. 2019).

2 Thermosensitive Drug Delivery Systems

Drug delivery systems sensitive to temperature have the ability to control the moment of drug delivery only when the human body temperature is above or below the physiological temperature.

Poly (N-isopropylacrylamide) (poly(NIPAAm)) is the most used thermoresponsive polymer in biomedical applications because it shows a sharp phase transition around 32 °C, which is close to the human body temperature (Fig. 1) (Heskins and Guillet 1968).

Fig. 1
figure 1

Poly(N-isopropylacrylamide) aqueous solution below and above the LCST

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The temperature at which this transition takes place is called the lower critical solution temperature (LCST). Below the LCST the polymer chains are surrounded by water molecules bound by hydrogen bonds, exposing an extended coil conformation and making the polymer soluble in water. However, when the temperature exceeds the LCST, the polymer loses the hydration water, becomes hydrophobic and precipitates (Fig. 1). Consequently, the cross-linked hydrogels obtained from these polymers swell under the LCST and shrink above the LCST (Fig. 2) (Fundueanu et al. 2009a).

Fig. 2
figure 2

Optical photomicrographs of poly(NIPAAm-co-acrylamide) microspheres in the swollen state in phosphate buffer below the LCST value at 22 °C (panel A) and above the LCST value at 45 °C (panel B). Reproduced with permission from Acta Biomaterialia (Fundueanu et al. 2009a)

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This swelling/shrinking process has been exploited for the development of thermosensitive drug delivery systems based on poly(NIPAAm-co-hydroxyethylacrylamide) microspheres (Fundueanu et al. 2013). The drug is usually released when the hydrogel is in the swollen state (below LCST). Above the LCST, the matrix shrinks and the release of drug is stopped. The pulsatile effect depends to the size of hydrogel microspheres. It was established that microspheres with the diameter ranging between 5 and 60 µm release the drug with almost the same rate below (in the swollen state) and above the LCST (in the collapsed state). On the contrary, microspheres with the diameter ranging between 125 and 220 µm release a significantly higher amount of indomethacin below than above the LCST. This difference was enough pronounced to ensure a pulsatile release mechanism when the temperature changes cyclically below and above the LCST.

In addition, the amount of entrapped drug influences the pulsatile release mechanism. Microspheres from poly(NIPAAm-co–N-vinylpyrrolidone) (poly(NIPAAm-co-NVP)) with a low loading degree (7.62%), display a substantially difference between the diclofenac (DF) release rate at temperatures situated below and above the LCST (Fig. 3A); this difference guarantees a pulsatile release mechanism (Fig. 3B). On the opposite, the release profiles of DF from microspheres with high loading degree (13.08%) at temperatures lower and higher the LCST are almost superimposable (Fig. 3C), therefore the cyclically variation of temperature does not ensure a pulsatile mechanism (Fig. 3D) (Fundueanu et al. 2020a).

Fig. 3
figure 3

Effect of temperature (panels A and B) and temperature cycling (panels C and D) (32 °C (empty symbols) and 38 °C (full symbols)) on DF release profiles from poly(NIPAAm-co-NVP) microspheres. The release studies were performed in PBS at pH = 7.4, using microspheres with 7.62% (w/w) DF (panel A and C) and with 13.08% (w/w) DF (panel B and D). Reproduced with permission from Express Polymer Letters (Fundueanu et al. 2020)

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As it was mentioned, the LCST of the poly(NIPAAm) in aqueous solution is around 32 ºC, which is slightly lower than the temperature of the human body. In order to increase the transition temperature, NIPAAm is copolymerized with hydrophilic co-monomers (Khan 2007; Liu et al. 2004). In fact, by copolymerization with hydrophilic monomers, the hydrophilic/hydrophobic balance is disturbed and higher energy (temperature) is necessary to break the hydrogen bonds between copolymer and water. In addition, the insertion of hydrophilic co-monomers along the main chain in a higher proportion alters the amide and isopropyl sequences of poly(NIPAAm) and can lead to a decrease or even loss of thermosensitive properties (Feil et al. 1933). On the opposite, the copolymerization of the NIPAAm with hydrophobic monomers leads to a decrease of the LCST below 32 °C (Luan et al. 2016). An atypical behaviour was observed for poly(N-isopropylmethacrylamide) (poly(NIPMAAm) with structure similar to that of NIPAAm with the exception of methyl groups. Although poly(NIPMAAm) is more hydrophobic than poly(NIPAAm), the phase transition temperature of poly(NIPMAAm) in aqueous solution is higher (≈46 °C) (Djokpé and Vogt 2001), reflecting the conformation of the monomeric structure in poly(NIPMAAm). The presence of methyl groups throughout the main polymeric chain hinders the hydrophobic interactions, therefore a higher temperature is necessary for polymer precipitation. The copolymerization of NIPAAm with increasing amount of NIPMAAm gives copolymers with increased values of LCST (Table 1) (Fundueanu et al. 2016).

Table 1 Dependence of LCST on the co-monomer ratio in the feed and in the copolymer. The concentration of the copolymer solution was 1% (w/v). Reproduced with permission from Macromolecular Chemistry and Physics (Fundueanu et al. 2016)
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Practically, the LCST of poly(NIPAAm-co-NIPMAAm) increases almost linearly with the percentage of NIPMAAm in the copolymer; comparable results were found by Djokpé and Vogt (Djokpé and Vogt 2001).

In fact, the phase transition of the copolymer is situated between the LCST of poly(NIPAAm) homopolymer and that of poly(NIPMAAm). Under simulated physiological conditions (phosphate buffer at pH 7.4) poly(NIPAAm) displays a LCST of 29.9 °C while poly(NIPMAAm) has a LCST at 44.8 °C, which both are lower than that determined in aqueous solution. This decrease in the transition temperature is due to the high ionic strength of the phosphate buffer which weakens the interaction between the polymer chains and the water.

However, the most used pharmaceutical form for these polymers in biomedical applications are hydrogels (Andrei et al. 2016; Sung et al. 2015). They are obtained by copolymerization of corresponding co-monomers in the presence of small amounts of cross-linkers. The resulted hydrogels should preserve the thermosensitive properties of the linear polymer. Moreover, they should display fast swelling deswelling rates to small changes of the temperature to ensure a pulsatile drug delivery (Fig. 4) (Fundueanu et al. 2016).

Fig. 4
figure 4

Influence of cyclically temperature change on dexamethasone release from the poly(NIPAAm-co-NIPMAAm) microgels. Data were obtained in simulated physiological conditions (PBS at pH = 7.4) at 32 °C (empty symbols) and 38 °C (full symbols). Reproduced with permission from Macromolecular Chemistry and Physics (Fundueanu et al. 2016)

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Swelling/desweling rates of thermosensitive hydrogels depend mainly by the diffusion rate of water molecules into and out of the hydrogel. The diffusion rate of water molecules is governed by the dimension and porosity of the hydrogel. Certainly, the smaller is the size the faster is the diffusion rate; in addition a high porosity allows a rapid diffusion of water (Sato-Matsuo and Tanaka 1988; Tanaka and Fillmore 1979). The creation of a porous structure in a three-dimensional network is one of the best procedure used to avoid the „skin effect” that hinders the diffusion of water and promotes a rapid volume change of hydrogel in response to temperature modifications (Strachotova et al. 2007). In hydrogels with porous structures, the water is ejected or absorbed by convection, a much faster process than by conventional diffusion. Usually, porous poly(NIPAAm) hydrogels were prepared in the presence of a template molecule. Tokuyama and Kanehara (2007) have synthesized hydrogels by free radical copolymerization of NIPAAm and N,N-methylene-bis-acrylamide, as cross-linker, in an oil-in-water emulsion (Tokuyama and Kanehara 2007). At the end of polymerization, the dispersed oil minidroplets used as template are removed by washing with appropriate solvents resulting a porous hydrogel. Thermally reversible macroporous poly(NIPAAm) hydrogel has also been synthesized in aqueous solution at a temperature above the LCST. As long as the reaction takes place at a temperature above the LCST, the polymer collapses as it reaches a certain degree of polymerization. The hydrophobic interactions that take place later between the aggregates create bridges that generate a heterogenous macroporous structure (Yan and Hoffman 1995). Hydrogels with large pores were also obtained in the presence of hydroxypropyl cellulose as a pore-forming agent. The response rate to the temperature changes was much faster for these hydrogels than those obtained without porogens. However, since the polymerization occurred above LCST, the macroporous hydrogel displayed a heterogeneous structure (Wu et al. 1992). Microgels with homogeneous pore size and size distribution were obtained by chemical cross-linking of preformed poly(NIPAAm-co-AAm) in the presence of poly(NIPAAm) thermosensitive oligomers, taken as porogen (Fundueanu et al. 2009b). The porous microspheres were obtained both below and above the LCST of the poly(NIPAAm) oligomer. The size of the pores is obviously influenced by the amount of the porogen in the reaction mixture. In the absence of the porogen no pores are formed and the microspheres show a compact internal structure and a smooth surface. The best size distribution and homogeneity of the pores were obtained under the LCST of porogen because the synthesis process takes place in a homogeneous solution of the polymer and oligomer (Fig. 5).

Fig. 5
figure 5

SEM micrographs of porous poly(NIPAAm-co-AAm) microspheres obtained using poly(NIPAAm) oligomers as porogens: cross-section (panel A) and surface detail (panel B). The bars correspond to 100 and 20 µm in panels A and B, respectively. Reproduced with permission from International Journal of Pharmaceutics (Fundueanu et al. 2009b)

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Due to the porous structure and hydrophilic character of the co-monomers, the microgels display relatively high values of water regain and swelling degree. Also, they are characterized by a very rapid response rate when the temperature changes below and above the body temperature. The porous microspheres have proven to be suitable matrix for loading and temperature-controlled release of the high molecular weight model drug blue dextran (Fig. 6).

Fig. 6
figure 6

Effect of temperature cycling (33 °C (empty symbols) and 38 °C (full symbols)) on Blue Dextran release from poly(NIPAAm-co-AAm) microspheres in phosphate buffer at pH 7.4. Reproduced with permission from International Journal of Pharmaceutics (Fundueanu et al. 2009b)

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3 PH-Sensitive Drug Delivery Systems

Drug delivery systems sensitive to pH were designed and developed to exploit the pH variations in the human body. For example, the pH varies from 1–2 in the stomach to 7.4 and even 8 in the blood and duodenum, respectively. Additionally, the pH-sensitive systems have appealed an enormous interest because the differences in pH between normal tissue and tumoral tissues can be used to design pH-sensitive drug delivery carriers that can target tumors and release loaded drugs at the tumor site (Xu et al. 2020; Men et al. 2020; Sim et al. 2019).

Generally, the pH-sensitive polymers comprise weakly acidic (carboxylic) or weakly basic (amine) functional groups (Riyajan 2019; Liu et al. 2018). The most important properties of these polymers result from the protonation/deprotonation of the pH-sensitive groups to small pH change in the range of physiological conditions. The linear polymers change their hydrophilic / hydrophobic character and consequently their solubility. The cross-linked hydrogels synthesized from these polymers swell and shrink as a result of protonation/deprotonation of functional groups (Vaghani and Patel 2011; Sabzi et al. 2020). Basically, hydrogels possessing carboxylic pH-sensitive groups become protonated at low pH, below the pKa and therefore is relatively not swollen (Fig. 7A). As the pH rises above the pKa value, the carboxylic groups ionize, become hydrophilic and the hydrogel swells. Hydrogels with amino pH-sensitive groups have a completely opposite behavior. At low pH, under the pKa, the amino groups are protonated (ionized) and the hydrogel swells while at higher pH values, above pKa, the amino groups are non-ionized, less hydrophilic and the hydrogel is in a collapsed state (Fig. 7B).

Fig. 7
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Swelling behavior of an anionic (panel A) and cationic hydrogel (panel B) in acidic and basic solution

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The most commonly used pH-sensitive polymers in the linear form are methacrylic acid copolymers, known as Eudragit R (a registered trademark of Rohm Pharmaceuticals; Darmstadt, Germany). For example, Eudragit L 100-55 T forms salts and dissolves above pH 5.5 and is insoluble in gastric fluids (Fig. 8). This property has been exploited to deliver drugs to the large intestine by way of pH sensitive enteric coatings or as micro- and nano-particles (Hao et al. 2013).

Fig. 8
figure 8

Chemical structure of Eudragit L 100–55

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Besides Eudragits, other pH-sensitive polymers such as derivatives of cellulose are largely used for drug targeting to the small intestine. Among these, cellulose acetate phthalate (pKa ̴ 5.5), also known as cellacefate is a cellulose derivative where about half of the hydroxyl groups are esterified with acetyls and a quarter is esterified with one or two carboxyls of the phthalic acid (Fig. 9). It is a commonly used polymer in the formulation of pharmaceuticals, such as the enteric coating of tablets or capsules for controlled release formulations (Jagdale and Chandekar 2017).

Fig. 9
figure 9

Chemical structure of cellulose acetate phthalate

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However, the most important drawback of these polymers is the lack of biodegradability. This inconvenience is not so severe for pharmaceutical formulations with oral application, instead, for other types of administrations such as parenteral ones, the biodegradability becomes a very acute problem. Therefore, natural polymers with native pH-sensitive units were frequently used for delivery of drugs modulated by the pH. Among these, alginic acid (Boi et al. 2020) and chitosan (Ata et al. 2020) are the most representative anionic and cationic natural polymers, respectively, used in biomedical applications. Concerning the chemical structure, alginic acid is described as a linear block copolymer composed of sequences with consecutive β-1,4-D-mannuronic acid residues (M-blocks), α-L-guluronic acid residues (G-blocks) and alternating M and G residues (MG-blocks) (Fig. 10).

Fig. 10
figure 10

Chemical structure of alginic acid

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The pKa of mannuronic and guluronic acid in 0.1 M NaCl are known to be 3.38 and 3.65, respectively. Owing to its pH-sensitivity, biocompatibility, low toxicity and muco-adhesive properties, alginic acid is considered a noticeable polymer for biomedical applications. Therefore, George and Abraham (2007) developed a pH-sensitive alginate–guar gum hydrogel cross-linked with glutaraldehyde for oral administration of protein drugs (George and Abraham 2007). It is well-known that protein instability mainly in gastric fluids is one of the major motives by which proteins are administered usually through injection rather than taken by oral route. Due to the protonation of alginate, the amount of released protein from test hydrogels was minimal at pH 1.2 (∼20%), but it was found to be significantly higher (∼90%) at pH 7.4., when the carboxylic groups are in ionized form and the hydrogel swells.

With the aim to enhance the drug entrapment efficiency and control the swelling properties and stability of drug delivery system in different pH media, Ca2+ cross-linking and freeze-thawing cycle techniques were used to prepare sodium alginate/poly(vinyl alcohol) (PVA) hydrogel beads (Hua et al. 2010). The alginate/PVA mixture was firstly cross-linked with Ca2+ to form beads and then subjected to freezing–thawing cycles for additional crosslinking. It was proved that the swelling, drug release profiles, and degradation of the developed beads were influenced by pH of the testing medium and PVA content.

Chitosan is another linear polysaccharide composed of randomly distributed β–linked D-glucosamine and N-acetyl-D-glucosamine (Fig. 11).

Fig. 11
figure 11

Chemical structure of chitosan showing the repeating subunits

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Like alginate, it is biocompatible, biodegradable and the degradation by-products are not toxic. It is the second abundant biomass-derived polysaccharide and is obtained by treating the chitin shells of shrimp and other crustaceans with sodium hydroxide. If alginate-based hydrogels collapse in an acidic environment and swell in a weakly basic environment, chitosan hydrogels have an opposite behaviour; they swell in acidic pH and collapse in basic media. This comportment is given by the presence of abundant amino groups with a pKa of around 6.5. Since the pKa value is situated in the pH range of tumor environment, it is widely used for controlled delivery of anticancer drugs (Ghaffari et al. 2020; Wang et al. 2019). As follows, an interesting hydrogel was designed and developed from the cross-linking of chitosan, graphene, and cellulose nanowhisker via Schiff base reaction by a synthetic dialdehyde (Omidi et al. 2020). The hydrogel was loaded with two antitumoral drugs, doxorubicin and curcumin.

The drug release profiles demonstrated that the rate of release is dependent on switching pHs from neutral pH to the acidic conditions. By decreasing of pH from 7.4–5.4, the strength of cross-linking imine bonds decreases and as well the repulsive ionic interactions between the protonated amino groups increase which leads to an expanded structure.

Chitosan was also used for the preparation of microcapsules for cell encapsulation and cell culture (Fundueanu et al. 2020b). Since the encapsulation process must take place in mild conditions and the neutral chitosan is soluble in just in acidic conditions, harmful to cells, the amino groups were protonated and the chitosan was transformed in the hydrochloride form. Thus, the chitosan derivative was easily solubilized in isotonic solutions and cell encapsulation was successfully performed with no cell damage. Moreover, the protonated chitosan microcapsules allow the cells to adhere to their surface.

One of the most important drawbacks of hydrogels prepared from simple alginate or chitosan is their low stability in physiological fluids. Therefore, core–shell nanoparticles from polyelectrolytes complexes between chitosan and alginate were prepared for oral insulin administration (Mukhopadhyay et al. 2015).

The nanoparticles displayed an almost spherical shape, an average diameter of 100–200 nm determined by dynamic light scattering, and high insulin encapsulation efficiency. In simulated gastric fluid, the nanoparticles retained almost entire amount of encapsulated insulin while in simulated intestinal conditions a sustained release was achieved. Additionally, in vivo experiments presented substantial intestinal absorption of insulin, showing noticeable hypoglycaemic effects with enhanced insulin bioavailability in diabetic mice. After oral administration, the core–shell nanoparticles displayed no systemic toxicity.

4 pH/Temperature-Sensitive Drug Delivery Systems

pH/temperature sensitive drug delivery systems were designed and developed to exploit the variation of both pH and temperature of the human body. They are obtained either in the form of linear or cross-linked polymers by copolymerization of NIPAAm with pH-sensitive monomers. The pH-sensitive co-monomer can increase or decrease the LCST if they are in the un-protonated or protonated state (Yoo et al. 2000). Also, the modification of pH and temperature can modulate the solubility, if the polymer is in the linear form, or swelling behavior for cross-linked hydrogels.

Thus, stimuli-sensitive statistical terpolymers of N-isopropylacrylamide (temperature-sensitive), butyl methacrylate (BMA) and acrylic acid (AA) (pH-sensitive) of various molecular weights were used to prepare microcapsules for modulating the release of insulin (Ramkissoon-Ganorkar et al. 1999). This system has the advantage that only physical interactions polymer/polymer and polymer/protein were used to prepare insulin–loaded microcapsules in mild conditions.

The release rate of insulin was controlled by the pH, temperature and molecular weights of polymers. In gastric fluid at pH = 1.2 and body temperature, the beads were not soluble, and therefore no drug is released in the stomach. At pH 7.4 and body temperature, the low molecular weight polymeric beads shown a dump-like profile and dissolved within 2 h (release mechanism controlled by bead dissolution), while the high molecular weight polymeric beads swelled only and released insulin gradually over a period of 8 h (release mechanism controlled by bead swelling and insulin diffusion).

In another work, the escape of DNA loaded poly(vinyl alcohol) microspheres from multinucleated microcapsules took place through the holes induced by dissolution of the pH/temperature sensitive polymer at colonic pH and temperature (Fundueanu et al. 2007).

Fundueanu et al. (2009) developed a pH/ temperature sensitive copolymer based on NIPAAm, methyl methacrylate (MM) and methacrilic acid (MA) (Fig. 12) (Fundueanu et al. 2009).

Fig. 12.
figure 12

Chemical structure of poly(NIPAAM-co-MM-MA)

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Poly (NIPAAm-co-MM-co-MA) with a co-monomer molar ratio of 79:13:8 is insoluble in the gastric fluid (pH = 1.2), but soluble in the intestinal fluid (pH = 6.8 and 7.4), at the body temperature (37 °C). This copolymer was mixed with the hydrophobic cellulose acetate butyrate (CAB) and vitamin B12 (taken as a water soluble drug model system) and transformed in spherical microparticles by oil-in-water solvent evaporation method. With the aim to determine the compatibility between the two components of the microspheres (CAB with pH/temperature sensitive copolymer), the glass transition temperatures (Tg) of separated polymers and of microspheres were determined.

The occurrence of two Tg values validated that there is no interaction between the two components; they precipitated separately during microsphere preparation, the smart polymer forming small micro-domains on the surface and within the microspheres (Fig. 13A).

Fig. 13
figure 13

SEM micrographs of CAB/poly(NIPAAm-co-MM-co-MA) microspheres before (panel A) and after (panel B) release studies in PBS at pH = 7.4. Reproduced with permission from Journal of Materials Science-Materials in Medicine (Fundueanu et al. 2009)

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At pH = 1.2 and 37 °C, poly(NIPAAm-co-MM-co-MA) is not soluble, therefore the amount of vitamin B12 released is very low. When the temperature is lowered to 20 °C, below LCST, the copolymer solubilizes and an increased percentage of vitamin B12 is released. In phosphate buffer at pH = 6.8, the copolymer is less soluble than at pH = 7.4, therefore the release rate is lower but is much higher than at pH = 1.2. Obviously, the highest amount of vitamin is released at pH = 7.4, when the copolymeric domains solubilize totally creating pores within three-dimensional network (Fig. 13B).

The copolymerization of NIPAAm with ionic co-monomers can increase or decrease the LCST if the charged co-monomers are in the protonated/unprotonated state (Yoo et al. 2000). However, the charged co-monomers, due to their high hydrophilicity in the ionized state, diminish dramatically or even annihilate the thermosensitivity of the copolymer (Constantin et al. 2014).

Remarkably, when the pH-sensitive units interact electrostatically with hydrophobic opposite charged bioactive compounds like diclofenac (Fundueanu et al. 2013), propranolol (Constantin et al. 2014), diphenhydramine (Fundueanu et al. 2017), indomethacin (Constantin et al. 2020), the copolymers restore the thermosensitive properties. Correspondingly, the hydrogels synthesized from these copolymers are in the “inactivated” state in the absence of specific biomolecules at normal physiological pH and temperature (pH = 7.4 and T = 36 °C). However, in the presence of bioactive biomolecules, hydrogels undergo “activation”, shrink and expel mechanically a certain amount of drug (Fig. 14).

Fig. 14
figure 14

Schematic representation of the principle of operation of pH/thermosensitive poly(NIPAAm-co-MA) microgels in the presence of the triggering agent, under simulated physiological conditions. Reproduced with permission from European Journal of Pharmaceutical Sciences (Constantin et al. 2014)

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It must be mentioned that the pH-sensitive component plays the role of a biosensor, the biomolecule acts as a triggering agent, and the poly(NIPAAm) represents the delivery component (actuator).

The proposed hydrogels represent the support of the next generation of self-regulated drug delivery systems based on a sensor (pH-sensitive units) able to detect the perturbation of physiological conditions and a delivery component (thermosensitive units) able to push the necessary dose of drug.

5 Conclusions

While the basic concepts of the stimuli-sensitive hydrogels are sound, the practical applications necessitate significant improvements in the hydrogel properties. For example, hydrogels do not have the mechanical strength indispensable in many applications. In addition, the response rate to the input signal is low for large-scale hydrogels. Difficulties to achieve pH/thermo-sensitivity within a narrow pH or temperature range (from pH 7.4 to pH = 7.0 or 6.5; from T = 36.5 °C to 37 °C) often appear since the transition temperature of the cross-linked hydrogels is not as abrupt as for linear polymers. Drug loading and drug retention within microgels for a long period of time could be a major drawback. Also, the selectivity of these systems must be substantially improved.