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cyclodextrin injectable hydrogel for the sustained release of drugs
Journal Pre-proofs A Thermoresponsive Hydrophobically Modified Hydroxypropylmethyl Cellulose/Cyclodextrin Injectable Hydrogel for the Sustained Release of Drugs Masanori Okubo, Daisuke Iohara, Makoto Anraku, Taishi Higashi, Kaneto Uekama, Fumitoshi Hirayama PII: DOI: Reference:
S0378-5173(19)30890-7 https://doi.org/10.1016/j.ijpharm.2019.118845 IJP 118845
To appear in:
International Journal of Pharmaceutics
Received Date: Revised Date: Accepted Date:
12 August 2019 18 October 2019 1 November 2019
Please cite this article as: M. Okubo, D. Iohara, M. Anraku, T. Higashi, K. Uekama, F. Hirayama, A Thermoresponsive Hydrophobically Modified Hydroxypropylmethyl Cellulose/Cyclodextrin Injectable Hydrogel for the Sustained Release of Drugs, International Journal of Pharmaceutics (2019), doi: https://doi.org/10.1016/ j.ijpharm.2019.118845
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© 2019 Published by Elsevier B.V.
A Thermoresponsive Hydrophobically Modified Hydroxypropylmethyl
Cellulose/Cyclodextrin
Injectable Hydrogel for the Sustained Release of Drugs Masanori Okuboa, Daisuke Ioharaa,b*, Makoto Anrakua,b, Taishi Higashic, Kaneto Uekamaa, Fumitoshi Hirayamaa
aFaculty
of Pharmaceutical Sciences and bDDS Research Institute, Sojo University, 4-22-1
Ikeda, Nishi-ku, Kumamoto 860-0082, Japan cGraduate
School of Pharmaceutical Science, Kumamoto University, Oe-honmachi,
Kumamoto 862-0973, Japan
*Corresponding Author: Daisuke Iohara E-mail: [email protected] Tel: +81-96-326-5084 Fax: +81-96-326-5048
Abbreviations : Cyclodextrin (CD), Hydroxypropylmethyl cellulose (HPMC), Hydrophobically modified hydroxypropylmethyl cellulose (HM-HPMC), Degree of substitution (DS), Indocyanine green (ICG), Phosphate buffered saline (PBS), Maximum drug concentration (Cmax), Time required to reach maximum drug concentration (Tmax), Area under the concentration-time curve from time 0 to infinity (AUC0-∞), Mean residence time from time 0 to infinity (MRT0-∞), Storage modulus (G’), Loss modulus (G’’).
ABSTRACT The objective of this study was to develop a thermoresponsive injectable hydrogel for the sustained release of drugs by taking advantage of host-guest interactions between a hydrophobically modified hydroxypropylmethyl cellulose (HM-HPMC) and cyclodextrin (CD).
A thermoresponsive injectable hydrogel was prepared by simply adding CDs to HM-
HPMC hydrogel. The HM-HPMC hydrogel was converted into a sol with a low viscosity through host-guest interactions with CDs. The HM-HPMC/β-CD hydrogel became a gel near body temperature where the host dissociated from the hydrophobic moieties of the polymer in response to the temperature. The yield stress of the HM-HPMC became progressively lower on the addition of β-CD which was desirable in the case of developing an injectable formulation. When the HM-HPMC/-CD hydrogel containing indocyanine green (ICG) was subcutaneously administered to mice, the fluorescence of the ICG remained relatively constant for 24 h after the administration, which was substantially longer than that for ICG alone or an HPMC formulation. The plasma insulin level was maintained for a longer period of time when the HM-HPMC/-CD containing insulin was administered and the MRT value was increased by 1.6 times compared to a solution of insulin alone. In addition, the HM-HPMC/β-CD hydrogel formulation showed a prolonged hypoglycemic effect in response to the insulin which was slowly released from the hydrogel. A thermoresponsive injectable hydrogel was successfully constructed from the highly viscous HM-HPMC and β-CD, and the resulting formulation functioned as a sustained release carrier for drugs.
KEY WORDS cyclodextrin ･ thermoresponsive ･ injectable ･ hydrophobically modified polymer ･ hydrogel
2
1. Introduction In the past few years, numerous biotechnology-based medicines have been approved for treating various diseases (Morrison and Lahteenmaki, 2015). However, there are still many challenges in terms of developing the ideal formulation of a biotechnology-based medicine such as proteins because of their inherent characteristics related to their chemical and physical stability, immunogenicity and pharmacokinetics (Frokjaer and Otzen, 2005; Wu and Jin, 2008). Therefore, many biotechnology-based medicines except for antibody require daily or multiple injections for years or even an entire lifetime, which can lead to patient noncompliance. It is therefore very important to address such issues in the formulation design of the protein drugs as well as the low molecular weight drugs especially for the treatment of chronic diseases such as malignant tumors, diabetes, hepatitis, leukemia, rheumatoid arthritis etc. Hydrogels are composed of three dimensional networks of hydrophilic polymers in which the polymer chains are chemically or physically inter or intra molecularly cross-linked. Hydrogels have been frequently used as drug or cell delivery vehicles and have recently attracted considerable attention for use in tissue engineering because they possess high water contents and are able to mimic extracellular matrixes (Thambi et al., 2017; Nagahama et al., 2018; Lee and Mooney, 2001; Vashist et al., 2014; Gaharwar et al., 2014). In comparison with covalently linked hydrogels, physically linked hydrogels that can be prepared by inter or intra molecular interaction such as hydrophobic or related interactions, hydrogen bonding and host– guest interactions (Delplace et al., 2016), show sol-gel transitions in response to various external stimuli (pH, light, temperature and oxidizing/reducing agent etc.) (Sood et al., 2016; Harada et al., 2014; Qiu and Park, 2001; Yan et al., 2012). Among these stimuli-responsive hydrogels, thermoresponsive hydrogels are a classical example of a clinically approved formulation for use as a drug delivery vehicle (Takeuchi et al., 1999) because heating is an external stimulus which is readily available in every part of the body after administration. These 3
hydrogels are typically form sols at room temperature that have a low viscosity, but form gels when they are warmed to body temperature. Such injectable hydrogels are of great interest as potential materials for prolonging the retention of a drug at the administered site, leading to the modification of the release profile of the drug or absorption rate of the drug into body (Kempe and Maeder, 2012; Li et al., 2012; Yu and Ding, 2008; Alexander et al., 2013). Block copolymers such as poly(N-isopropylacrylamide) (Jeong et al., 2002; Soppimath et al., 2002; Alexander et al., 2014) or poly (ethylene oxide) (PEO)/poly(propylene oxide) (PPO) tri-block copolymers (Ruel-Gariepy and Leroux, 2004; Zhao et al., 2016; Park et al., 2009; Alexander et al., 2016) were, in the past, examined for this purpose, because such polymers undergo phase separation that is dependent on the environmental temperature (Parmar et al., 2017). However, higher polymer concentrations (above 20 w/v %) (Hoare and Kohane, 2008) was generally required for such a thermal gelation, which is costly and can cause allergic reactions. We recently developed a new strategy for preparing such thermoresponsive materials based on interactions between cyclodextrin (CD) and hydrophobically modified polymers (Iohara et al., 2017). Generally, the inclusion process for a CD is dependent on the temperature being used and is frequently occurs at low temperatures, while heating results in dissociation (Rekharsky and Inoue, 1998, 2000; Brocos et al., 2010). By simply adding CD to a hydrophobically modified hydroxypropylmethyl cellulose (HM-HPMC) hydrogel the gel became a low viscous sol as the result of the host-guest interaction of CD with the hydrophobic moiety. On heating, the CDs dissociate from the HM-HPMC/CD, resulting in the sol state turning into a gel state at around body temperature. On the basis of the previous study, we report here on the development of a thermoresponsive injectable hydrogel for the purpose of the sustained release of drugs. The HM-HPMC/CD is a sol with a low viscosity at low temperatures, which allows it to be easily injected, whereas it turns into a high viscous gel at the injection site (Fig. 1). An appropriate formulation of an injectable hydrogel was determined 4
by focusing on the thermoresponsive sol-gel transition property, and the release profile of insulin, a model drug, from the hydrogel was assessed after subcutaneous administration of the hydrogel.
2. Materials and method
2.1.
Materials
HM-HPMC (commercially known as Sangelose 60L or Sangelose 90L) and HPMC (METOLOSE 90SH - 30000) were obtained from the Daido Chemical Co. (Osaka, Japan). 60L HM-HPMC (M.W. 400,000) and 90L HM-HPMC (M.W. 600,000) contain a C-18 fatty acid group at the hydroxypropyl ends and has a degree of substitution (DS) of 0.3-0.6 mol %. The DS of the hydroxypropoxy and methoxy groups in 60L HM-HPMC were 7.0-11.0 mol % and 27.0-30.0 mol%, respectively. The 90L HM-HPMC contains hydroxypropoxy and methoxy groups 7.0-11.0 mol% and 21.5-24.0 mol%, respectively. The α-, - and -CD were obtained from Nihon Syokuhin Kako Co. (Tokyo, Japan). Indocyanine green (ICG) was purchased from Daiichi Sankyo Co. (Tokyo, Japan). Human insulin was purchased from Sigma Aldrich Co. (Tokyo, Japan). All other reagents were of analytical grade.
2.2.
Preparation of Thermoresponsive HM-HPMC/CD Injectable Hydrogels
HM-HPMC was added to a heated phosphate buffered saline (PBS) solution and dispersed by agitating. On cooling, the suspension of HM-HPMC was isolated as a clear HM-HPMC 5
hydrogel. Thermoresponsive HM-HPMC/CD hydrogels were prepared by simply adding various amount of α-, - and -CDs in the HM-HPMC hydrogels. The resulting HMHPMC/CD hydrogels were characterized by following rheological studies.
2.3.
Rheological Studies
Rheological investigations of the HM-HPMC/CD hydrogels were performed with a MCR-101 rheometer (Anton Paar Japan K.K., Tokyo, Japan). A cone and plate geometry with a diameter of 25 mm was used for these measurements. To evaluate the effect of CDs on the viscosity of HM-HPMC, steady shear rheology measurements were performed at 25.0 ºC with a shear rate at 0.1 s-1. Thermoresponsive property of HM-HPMC/CD hydrogels was evaluated by the viscosity measurement on the dynamic temperature and the dynamic rheology measurements. The viscosity of HM-HPMC/CD hydrogels was determined at a heating rate of 3.0 ºC/min at a shear rate of 1.0 s-1. The temperature of samples for the dynamic rheology was controlled to 10, 20, 30 and 37 ± 0.01 º C and the dynamic frequency data were obtained in the linear viscoelastic regime. For the usability test of the hydrogel for an injection, the measured strain was plotted as a function of shear stress.
2.4.
In Vivo Release Profile of ICG from the HM-HPMC/β-CD Hydrogel
A solution of the ICG was mixed with the HPMC or HM-HPMC/-CD hydrogel, and the final concentration of each component was fixed at 0.5 M for ICG, 0.5 w/v % for HPMC or HMHPMC and 0.04 w/v % for -CD, respectively. The HM-HPMC/-CD hydrogel containing 6
ICG was administered subcutaneously to the dorsal skin of ddy mice (male, 7 weeks of age) using a 1 ml syringe with a 26 gauge needle. The fluorescence intensity of ICG was measured using a fluorescence imaging apparatus (IVIS Lumina XR, PerkinElmeer Co., Waltham, MA, USA) at 0, 6, 12 and 24 h.
2.5.
Evaluation of HM-HPMC/β-CD Hydrogel for a Sustained Release of Insulin
A sample of the HM-HPMC (0.5 w/v %)/-CD (0.04 w/v %) product containing 3 U/kg of insulin was prepared for their administration. Insulin solution was simply mixed with HMHPMC/β-CD and the insulin did not affect the viscosity of the hydrogel. The hydrogel was administered subcutaneously to the dorsal skin of rats (male, 7 weeks of age) using a 1 ml syringe with a 26 gauge needle. Blood samples were collected from the tail vein at predetermined interval after the administration. Plasma insulin and glucose levels in the samples were measured using an Insulin ELISA KIT(TMB) (AKRIN-010T, Shibayagi, Gunma, Japan) or a Glucose-CII-Test Wako (Tokyo, Japan), respectively, following the manufacturer’s recommended protocol. A non-compartment model was used for the pharmacokinetic analysis. Pharmacokinetic parameters, including maximum drug concentration (Cmax), time required to reach maximum drug concentration (Tmax), area under the concentration-time curve from time 0 to infinity (AUC0- ∞ ) and mean residence time from time 0 to infinity (MRT0- ∞ ) were calculated by the moment analysis program which is available on Microsoft Excel.
2.6.
Statistical Analysis
7
Data are presented as the mean values from n samples, and the results are reported as the mean ± S.E. Kruskal-Wallis analysis of variance (ANOVA) was performed followed by student’s ttests for comparison of samples. For all analyses, values of p < 0.05 were regarded as statistically significant.
3. Result and discussion
3.1. Preparation and Evaluation of Thermoresponsive HM-HPMC/CD Injectable Hydrogel for Injection
3.1.1. Modification of the Viscosity of HM-HPMC Hydrogels by CDs HM-HPMC samples with different molecular weights (60L HM-HPMC and 90L HMHPMC) were tested as the base of the injectable hydrogel. HM-HPMC solutions are highly viscous due to the intra- and inter-molecular associations of the stearyl groups which are introduced at the hydroxypropyl ends (Ghosal et al., 2012; Ikeda et al., 1994). We previously reported that the high viscosity of a 60L HM-HPMC hydrogel was markedly decreased in the presence of small amounts of α-CD (Iohara et al., 2017). The small cavity of the α-CD allowed the stearyl moiety of the HM-HPMC to be selectively incorporated, which inhibited the associations of the polymers. The viscosity of the 60L and 90L HM-HPMC hydrogels was first compared in the presence of CDs in order to design an appropriate formulation for an injectable hydrogel. Figure 2 shows the viscosity of the 60L and 90L HM-HPMC hydrogels at a shear rate of 0.1 s-1. The high viscosity was observed for both the 60L and 90L HM-HPMC hydrogels alone at low concentration (0.5 w/v %), the values of which were about 34 Pa･s for 60L HMHPMC and 53 Pa･s for 90L HM-HPMC. Such a high viscosity of the 60L and 90L HM-HPMC 8
decreased by the addition α-CD or β-CD, and the reducing effect was dependent on the added amount of CDs. In the case of -CD, the reducing effect on the viscosity was small and a high viscosity was observed, even when 0.2 w/v% of -CD was added. In comparing the viscosity among the two HM-HPMC samples, the viscosity of the 90L HM-HPMC was higher than that of the 60L HM-HPMC over the entire range of CD concentrations. This can be attributed to the fact that the 90L HM-HPMC with a higher molecular weight contains a higher number of stearyl moieties (15 molecules per polymer chain) compared with 60L HM-HPMC (10 molecules per polymer chain). The viscosity of 90L HM-HPMC as well as 60L HM-HPMC was modified by the addition of α- and β-CD. A hydrogel with a higher viscosity would be expected to release drugs more slowly from its gel networks and for this reason, we chose to use the 90L HM-HPMC as the base for the injectable hydrogel.
3.1.2. Thermoresponsive Changes in Viscosity of the 90L HM-HPMC/CD Hydrogel Figure 3 shows the change in viscosity of the 90L HM-HPMC as a function of dynamic temperature. The high viscosity of the 90L HM-HPMC alone decreased with increasing temperature in the range of 20 to 45ºC similar to the behavior of a typical elastic polymer solution. In contrast, the viscosity of the 90L HM-HPMC/α-CD was low at low temperatures, but the viscosity increased slightly with temperature in the range of 40~45ºC in the presence of 0.04% α-CD. Such changes in viscosity were minor in the presence of 0.1% or 0.2% α-CD. Regarding the -CD, the viscosity clearly changed with increasing temperature, reaching a maximum viscosity at around 43ºC at 0.04 w/v% -CD. In the presence of -CD, a similar viscosity change was observed for 0.2 w/v% -CD and the maximum viscosity occurred at approximately 36ºC. We previously reported that the interaction of CD with a C-18 fatty acid is dependent on the size of the CD cavity; thus, the α-CD with a small cavity size showed a high affinity for the stearyl moiety, followed by the -CD and the -CD. In addition, the 9
interaction between the CD and the stearyl moiety varied with temperature and the dissociation appeared to occur at a higher temperature (Fig. S1). The different rheological behavior of the 90L HM-HPMC in the presence of the CDs can be explained by the fact that the viscosity of such systems is controlled by the host-guest interactions. The viscosity of the 90L HMHPMC/α-CD barely changed on heating because the stearyl moieties were contained within the α-CD which has a relatively higher affinity for stearyl moieties, even at higher temperatures. In the case of β-CD, a clear viscosity change was observed at 0.04% β-CD, where the dissociation of the β-CD occurred on heating. Because -CD has a low affinity for stearyl moieties, they readily dissociate, leading to a maximum viscosity at lower temperatures compared with the β-CD system. These rheological behaviors were different from the previously reported 60L HM-HPMC/CDs system (Fig. S1). The peak for the maximum was shifted to a higher temperature in the case of 90L HM-HPMC, probably because of the large number of stearyl moieties per molecule. The viscosity of the 90L HM-HPMC/-CD hydrogel was measured upon heating and cooling cycles (Fig. 4). The 90L HM-HPMC (0.5 w/v%)/βCD (0.04 w/v%) hydrogel was used for the measurement because it showed a maximum viscosity near body temperature. The viscosity of the 90L HM-HPMC/-CD hydrogel was low below 1.0 Pa･s at 10ºC, whereas it was six times as high at 37ºC. The changes in viscosity were markedly larger compared with those of the previous 60L HPMC/α-CD system (2.0 Pa‧ s at low temperature versus 4.0 Pa‧s at high temperature). The change was reversible upon heating and cooling because the inclusion process of the CD is in an equilibrium state. Temperature dependent viscosity changes were evaluated by dynamic rheology measurements (Fig. 5). The 90L HM-HPMC/-CD hydrogel showed a tan (loss modulus G’’/storage modulus G’) of > 1 at 10ºC, indicating that it was in a sol state. It should be noted that the tan value was ≒ 1 at 20ºC, indicating that it was in a sol-gel intermediate state. The tan values became < 1 at 30ºC and 37ºC, indicating that they were in a gel state. These collective results 10
suggest that the 90L HM-HPMC/-CD hydrogel was converted from a sol into a gel in response to a temperature change. A thermoresponsive sol-gel transition system could be constructed with 90L HM-HPMC by adding a CD with the appropriate cavity size and concentration.
3.1.3.
Rheological Properties of 90L HM-HPMC/CDs Hydrogel for an Injection
A hydrogel is generally a useful carrier for the sustained release of drugs, but the high viscosity of the hydrogel limits its use in practical applications. For use of a hydrogel in the formulation of an injectable formulation, the hydrogel must be readily ejected from the needle of the syringe. The rheological properties of the 90L HM-HPMC/CD hydrogels were evaluated in terms of the usability of the preparation for an injection. The strain of samples was plotted against shear stress, as shown in Fig. 6. Regarding the 90L HM-HPMC alone, the strain rapidly increased at 16 Pa indicating that a reasonable amount of force was required to collapse the gel structure and to initiate a flow. Such a sharp change in strain could be eliminated when the above system was used and a pattern similar to a PBS could be achieved. The value for shear stress at the inflection point (yield stress) was plotted against CDs concentrations in Fig 6(d). The yield stress of the 90L HM-HPMC hydrogel was decreased by adding CDs, and α-CD and β-CD showed a remarkable decrease in yield stress at low concentrations, i.e., when 0.02 w/v% α-CD or 0.05 w/v% -CD were added. Regarding the -CD, the effect on yield stress of the 90L HM-HPMC was weak and at least 0.5 w/v% was needed to decrease the yield stress. These results indicate that a 90L HM-HPMC/CD hydrogel with a low yield stress would be applicable to the injectable formulation. The viscosity of the 90L HM-HPMC/CDs hydrogels was also measured as a function of share rate (Fig. S2). The viscosity decreased with share rate showing quasi viscous flow properties. As expected, α-CD and -CD reduced the viscosity of 90L HMHPMC in the concentration dependent manner. On the basis of these collective data, we used 11
the 90L HM-HPMC (0.5 w/v%)/β-CD (0.04 w/v%) hydrogel for the formulation of the injectable hydrogel because the system showed a low yield stress and viscosity at low temperature as well as undergoing thermal gelation near body temperature (Fig. 3). α-CD was useful to modify the viscosity of the 90L HM-HPMC, however, the system did not undergo thermal gelation near body temperature. In the case of the γ-CD system, high CD concentrations were needed to modify the viscosity of the 90L HM-HPMC. Although the HMHPMC (0.5 w/v%)/γ-CD (0.2 w/v%) hydrogel underwent thermal gelation near body temperature, the yield stress and the viscosity of the formulation (Fig. 6 and Fig. S2) were not sufficiently adaptable for the formulation of the injectable hydrogel compared to the β-CD formulation. The 90L HM-HPMC (0.5 w/v%)/β-CD (0.04 w/v%) hydrogel could be readily ejected from the syringe with a 27G needle, whereas this was not possible for the HM-HPMC alone (Movie S1). The HM-HPMC (0.5 w/v%)/β-CD (0.04 w/v%) hydrogel, which undergoes thermal gelation near body temperature, was used as the base for the injectable hydrogel in the following in vivo studies.
3.2.
In Vivo Profile for the release of a Fluorescent Dye from the 90L HM-HPMC/β-CD
Hydrogel The release properties of the drug from the 90L HM-HPMC (0.5 % w/v)/-CD (0.04 % w/v) hydrogel was evaluated in in vivo studies. The 90L HM-HPMC/-CD hydrogel containing indocyanine green (ICG) was administered subcutaneously to mice and the fluorescence of the ICG was monitored at predetermined intervals (Fig. 7). The fluorescence of the ICG decreased dramatically within 12 h after administration and disappeared by 24 h in the case of the ICG solution alone. This result indicate that ICG rapidly diffused into the body from the solution. Similar behavior was observed for the HPMC formulation. On the other hand, a strong 12
fluorescence continued to be observed at 12 h after the administration of the 90L HM-HPMC/CD hydrogel, and fluorescence was also observed even in 24 h. Thermal gelation occurred as soon as the hydrogel was warmed to body temperature. Thus, it was estimated that the thermal gelation occurred within a few minutes after its administration because the injection volume (0.1 ml) was small enough to permit it to be heated rapidly. It should be noted here that the viscosity of the 90L HM-HPMC/β-CD hydrogel was essentially the same as that of HPMC at room temperature (Fig. S3). It can thus be concluded that the extended release of ICG from the 90L HM-HPMC/β-CD hydrogel can be attributed to the thermosresponsive sol-gel transition on being warmed to body temperature.
3.3. Evaluation of the Thermoresponsive 90L HM-HPMC/-CD Hydrogel for the Sustained Release of a Drug In order to evaluate the sustained-release property of the 90L HM-HPMC/-CD hydrogel, insulin (as a model drug) was loaded into the hydrogel. We confirmed that insulin had no effect on the rheological properties of the hydrogel (data not shown). Figure 8(a) shows the plasma insulin concentrations following the subcutaneous administration of 90L HM-HPMC/-CD hydrogel containing insulin. The maximum levels for plasma insulin were reached within 1 h after the administration of an insulin in an ordinary solution (Cmax: 42.8 ± 2.8 ng/ml), and the administered insulin disappeared from the blood circulation within 4 h. When the insulin in the 90L HM-HPMC/β-CD hydrogel formulation was administered, the insulin concentration also reached the maximum level within 1 h after administration (Cmax: 33.6 ± 1.7 ng/ml), but, in this case, the plasma insulin level at 4 h and 6 h after the administration was significantly higher than that of the insulin alone solution. The pharmacokinetics data for these formulations are shown in Table 1. The MRT value for the 90L HM-HPMC/β-CD hydrogel was prolonged by 13
about 1.6 times compared with the insulin alone solution (90L HM-HPMC/-CD = 2.6 ± 0.2 h, insulin alone solution = 1.6 ± 0.1 h). The plasma glucose levels were also monitored after the administration of 90L HM-HPMC/-CD hydrogel (Fig. 8(b)). When the insulin alone solution was administered to rats, the plasma glucose level dropped to about 40 mg/dl at 2 h after the administration, and the value recovered to normal levels within 6 h. In sharp contrast, the low plasma glucose level was maintained even at 6 h in the case of the administration of the 90L HM-HPMC/-CD hydrogel formulation (40 mg/dl at 6 h). The hypoglycemic effect was obviously prolonged for the 90L HM-HPMC/-CD hydrogel formulation. It was difficult to administer the insulin in the HM-HPMC formulation as shown in Movie S1. The HM-HPMC was converted into a sol with a low viscosity when the β-CD was added that permitted it to be administered from the syringe. The administered 90L HM-HPMC/β-CD hydrogel became a highly viscous gel by warming at body temperature, thus the insulin was slowly released from the gel matrix resulting in a prolonged hypoglycemic effect.
Table 1. In vivo pharmacokinetic parameters of insulin and the 90L HM-HPMC (0.5 w/v%)/CD (0.04 w/v%) hydrogel containing insulina AUC0-∞d
Sample
Cmaxb (ng/ml)
Tmaxc (h)
Insulin alone
42.8 ± 2.8
0.8 ± 0.2
79.0 ± 6.2
1.6 ± 0.1
33.6 ± 1.7*
0.5 ± 0.0
85.2 ± 3.5
2.6 ± 0.2*
Insulin/ 90L HM-HPMC/-CD aEach
(ng/ml･h)
value represents the mean ± S.E. of 3 experiments.
bMaximum
drug concentration.
cTime
required to reach maximum drug concentration.
dArea
under the curve from time 0 to infinity.
eMean
residence time from time 0 to infinity.
*p < 0.05 versus insulin. 14
MRT0-∞e (h)
4. Conclusions Thermoresponsive injectable hydrogels were successfully prepared by using a combination of the highly viscous 90L HM-HPMC and β-CD. The 90L HM-HPMC/β-CD underwent thermal gelation near body temperature as the result of host–guest interactions between the stearyl moieties of HM-HPMC and β-CD cavity. The 90L HM-HPMC/β-CD hydrogel was easy to administer by simple injection. The 90L HM-HPMC/β-CD turned into a high viscous gel after being entering the body, which allowed the drugs to be slowly released and efficacy to be maintained. The thermal response of the HM-HPMC/CD material can be tuned by selecting an appropriate CD and the concentrations of the components. A thermoresponsive injectable hydrogel, in which interactions between the CD and hydrophobically modified polymers determine its properties, promises to be useful as an advanced drug carrier for drugs incluiding proteins
Declaration of competing interest None.
Figure captions
Fig. 1.
Formulation of the sustained release system using a thermoresponsive HM-
HPMC/CD hydrogel.
Fig. 2. Changes in the viscosity of 60L and 90L HM-HPMC hydrogels (0.5 w/v%) by adding CDs. The shear viscosity at 0.1 s-1 was plotted as a function of CDs concentrations. Each point represents the mean ± S.E. of 3 experiments. 15
Fig. 3. Changes in the viscosity of the 90L HM-HPMC/CDs hydrogel (0.5 w/v%) on the dynamic temperature. The viscosity was determined at a shear rate of 1.0 s-1. Each point represents the mean ± S.E. of 3 experiments. Fig. 4. Thermoresponsive reversible viscosity change of the 90L HM-HPMC (0.5 w/v%)/CD (0.04 w/v%) hydrogel. The viscosity was determined at a shear rate of 1.0 s-1. Each point represents the mean ± S.E of 3 experiments.
Fig. 5. Loss tangents (tan ) for the 90L HM-HPMC (0.5 w/v%)/-CD (0.04 w/v%) hydrogel at 10, 20, 30 and 37℃. Each point represents the mean ± S.E of 3 experiments. ※tan (loss tangent) = G’’ (loss modulus) / G’ (storage modulus)
Fig. 6. Stress-strain curves of the 90L HM-HPMC hydrogel (0.5 w/v%) in the presence and absence of α-CD (a), β-CD (b), γ-CD (c) and changes in yield stress as a function of CD concentrations (d). Each point represents the mean ± S.E. of 3 experiments. Fig. 7. In vivo fluorescence image after subcutaneous administration of an indocyanine green (ICG) solution, the HPMC hydrogel (0.5 w/v%) containing ICG and the 90L HM-HPMC (0.5 w/v%)/-CD hydrogel containing ICG.
Fig. 8. Plasma levels for insulin (a) and glucose (b) after the subcutaneous administration of a phosphate buffered solution, an insulin solution (3 U/kg) and the 90L HM-HPMC (0.5 w/v%)/-CD hydrogel (0.04 w/v%) containing insulin (3 U/kg) to rats. Each data point represents the mean ± S.E. of 3 experiments. *, p<0.05 versus insulin.
Acknowledgement
This work was supported, in part, by JSPS KAKENHI Grant Number 18K06618. The authors are grateful to Mr. Yamada, Ms. Take and Ms. Miyata for technical assistance.
16
References Alexander, A., Ajazuddin, Khan, J., Saraf, S., Saraf, S., 2013. Poly(ethylene glycol)poly(lactic-co-glycolic acid) based thermosensitive injectable hydrogels for biomedical applications. J. Controlled Release 172, 715-729. https://doi.org/10.1016/j.jconrel.2013.10.006. Alexander, A., Ajazuddin, Khan, J., Saraf, S., Saraf, S., 2014. Polyethylene glycol (PEG)Poly(N-isopropylacrylamide) (PNIPAAm) based thermosensitive injectable hydrogels for biomedical applications. Eur. J. Pharm. Biopharm. 88, 575-585. https://doi.org/10.1016/j.ejpb.2014.07.005. Alexander, A., Saraf, S., Saraf, S., 2016. Understanding the role of poloxamer 407 based thermoreversible in situ gelling hydrogel for delivery of PEGylated melphalan conjugate. Curr. Drug Delivery, 13, 621-630. https://doi.org/10.2174/1567201813666160204114000. Brocos, P., Diaz-Vergara, N., Banquy, X., Perez-Casas, S., Costas, M., Pineiro, A., 2010. Similarities and differences between cyclodextrin-sodium dodecyl sulfate host-guest complexes of different stoichiometries: molecular dynamics simulations at several temperatures. J. Phys. Chem. B 114, 12455-12467. https://doi.org/10.1021/jp103223u. Delplace, V., Obermeyer, J., Shoichet, M.S., 2016. Local Affinity Release. ACS Nano 10, 6433-6436. https://doi.org/10.1021/acsnano.6b04308. Frokjaer, S., Otzen, D.E., 2005. Protein drug stability: a formulation challenge. Nat. Rev. Drug Discovery 4, 298-306. https://doi.org/10.1038/nrd1695. Gaharwar, A.K., Peppas, N.A., Khademhosseini, A., 2014. Nanocomposite hydrogels for biomedical applications. Biotechnol. Bioeng. 111, 441-453. https://doi.org/10.1002/bit.25160. Ghosal, K., Chandra, A., Rajabalaya, R., Chakraborty, S., Nanda, A., 2012. Mathematical modeling of drug release profiles for modified hydrophobic HPMC based gels. Pharmazie 67, 147-155. https://doi.org/10.1691/ph.2012.1014. Harada, A., Takashima, Y., Nakahata, M., 2014. Supramolecular polymeric materials via cyclodextrin-guest interactions. Acc Chem. Res. 47, 2128-2140. https://doi.org/10.1021/ar500109h. Hoare, T.R., Kohane, D.S., 2008. Hydrogels in drug delivery: Progress and challenges. Polymer 49, 1993-2007. https://doi.org/10.1016/j.polymer.2008.01.027. Iohara, D., Okubo, M., Anraku, M., Uramatsu, S., Shimamoto, T., Uekama, K., Hirayama, F., 2017. Hydrophobically modified polymer/α-cyclodextrin thermoresponsive hydrogels for use in ocular drug delivery. Mol. Pharmaceutics 14, 2740-2748. https://doi.org/10.1021/acs.molpharmaceut.7b00291. Ikeda, K., Saitoh, I., Oguma, T., Takagishi, Y., 1994. Applications of hydrophobicallymodified hydroxypropyl methyl cellulose to topical preparations. III. Effect of 17
hydrophobically-modified hydroxypropyl methyl cellulose on the crystallization from supersaturated solutions of indomethacin. Chem. Pharm. Bull. 42, 2320-2326. https://doi.org/10.1248/cpb.42.2320. Jeong, B., Kim, S.W., Bae, Y.H., 2002. Thermosensitive sol-gel reversible hydrogels. Adv. Drug Delivery Rev. 54, 37-51. https://doi.org/10.1016/j.addr.2012.09.012. Kempe, S., Maeder, K., 2012. In situ forming implants - an attractive formulation principle for parenteral depot formulations. J. Controlled Release 161, 668-679. https://doi.org/10.1016/j.jconrel.2012.04.016. Lee, K.Y., Mooney, D.J., 2001. Hydrogels for tissue engineering. Chem. Rev. 101, 18691879. https://doi.org/10.1021/cr000108x. Li, Y., Rodrigues, J., Tomas, H., 2012. Injectable and biodegradable hydrogels: gelation, biodegradation and biomedical applications. Chem. Soc. Rev. 41, 2193-2221. Morrison, C., Lahteenmaki, R., 2015. Public biotech in 2014-the numbers. Nat. Biotechnol., 33, 703-709. https://doi.org/10.1039/C1CS15203C. Nagahama, K., Oyama, N., Ono, K., Hotta, A., Kawauchi, K., Nishikata, T., 2018. Nanocomposite injectable gels capable of self-replenishing regenerative extracellular microenvironments for in vivo tissue engineering. Biomater. Sci. 6, 550-561. https://doi.org/10.1039/C7BM01167A. Park, K.M., Lee, S.Y., Joung, Y.K., Na, J.S., Lee, M.C., Park, K.D. 2009. Thermosensitive chitosan-pluronic hydrogel as an injectable cell delivery carrier for cartilage regeneration. Acta Biomater. 5, 1956-1965. https://doi.org/10.1016/j.actbio.2009.01.040. Parmar, I.A., Shedge, A.S., Badiger, M.V., Wadgaonkar, P.P., Lele, A.K., 2017. Thermoreversible sol-gel transition of aqueous solutions of patchy polymers. RSC Adv. 7, 51015110. https://doi.org/10.1039/C6RA27030A. Qiu, Y., Park, K., 2001. Environment-sensitive hydrogels for drug delivery. Adv. Drug Delivery Rev. 53, 321-339. https://doi.org/10.1016/S0169-409X(01)00203-4. Rekharsky, M.V., Inoue, Y., 1998. Complexation thermodynamics of cyclodextrins. Chem. Rev. 98, 1875-1917. https://doi.org/10.1021/cr970015o. Rekharsky, M.V., Inoue, Y., 2000. Chiral recognition thermodynamics of β-cyclodextrin: the thermodynamic origin of enantioselectivity and the enthalpy-entropy compensation effect. J. Am. Chem. Soc. 122, 4418-4435. https://doi.org/10.1021/ja9921118. Ruel-Gariepy, E., Leroux, J. C., 2004. In situ-forming hydrogels-review of temperaturesensitive systems. Eur. J. Pharm. Biopharm. 58, 409-426. https://doi.org/10.1016/j.ejpb.2004.03.019. Sood, N., Bhardwaj, A., Mehta, S., Mehta, A., 2016. Stimuli-responsive hydrogels in drug delivery and tissue engineering. Drug Delivery 23, 748-770. https://doi.org/10.3109/10717544.2014.940091. Soppimath, K.S., Aminabhavi, T.M., Dave, A.M., Kumbar, S.G., Rudzinski, W.E., 2002. 18
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19
Fig. 1.
HM-HPMC R = -H -CH3 -CH2CH(OH)CH3
addition of CDs
HM-HPMC/CD
-CH2CH(OH)CH2OC18H37
heating OH
heating
O
HO
OH
O OH
O
O HO
HO
OH O HO
OH
O HO
O OH
OH OH
O OH
cooling
addition of CDs
OH
O OH
O
OH
heating
HO
heating
gel
sol
O
O
α-CD (n = 1) β-CD (n = 2) γ-CD (n = 3)
Subcutaneous injection
sol
Body temperature cooling (37℃)
gel
Easy to inject Sustained release of drug Fig. 1. Formulation of the sustained release system using a thermoresponsive HM-HPMC/CD hydrogel.
20
Fig. 2.
60
60
50
○：60L HM-HPMC ●：90L HM-HPMC
40
Viscosity (Pa・s)
Viscosity (Pa・s)
50
30 20
○：60L HM-HPMC ●：90L HM-HPMC
40 30 20 10
10 0
0 0
0.05
0.1
0.15
0.2
0
0.05
Concn. of α-CD (w/v%)
0.1
0.15
Concn. of β-CD (w/v%)
60
○：60L HM-HPMC ●：90L HM-HPMC
Viscosity (Pa・s)
50 40 30 20 10 0 0
0.05
0.1
0.15
0.2
Concn. of γ-CD (w/v%)
Fig. 2. Changes in the viscosity of 60L and 90L HM-HPMC hydrogels (0.5 w/v%) by adding CDs. The shear viscosity at 0.1 s-1 was plotted as a function of CDs concentrations. Each point represents the mean ± S.E. of 3 experiments.
21
0.2
Fig. 3.
Fig. 3. Changes in viscosity of the 90L HM-HPMC/CDs hydrogel (0.5 w/v%) on the dynamic temperature. The viscosity was determined at a shear rate of 1.0 s-1. Each point represents the mean ± S.E. of 3 experiments.
22
Fig. 4.
7
Viscosity (Pa･s)
6 5 4 3 2 1 0 10
37
10
37
10
37
10
37
10
Temperature (℃)
Fig. 4. Thermoresponsive reversible viscosity change of the 90L HM-HPMC (0.5 w/v%)/-CD (0.04 w/v%) hydrogel. The viscosity was determined at a shear rate of 1.0 s-1. Each point represents the mean ± S.E of 3 experiments.
23
Fig. 5.
Fig. 5. Loss tangents (tan ) for the 90L HM-HPMC (0.5 w/v%)/-CD (0.04 w/v%) hydrogel at 10, 20, 30 and 37℃. Each point represents the mean ± S.E of 3 experiments. tan (loss tangent) = G’’ (loss modulus) / G’ (storage modulus)
24
Fig. 6.
(b)
106
106
104
104
102
100 100
Strain (%)
Strain (%)
(a)
102
● : HM-HPMC ● : w ith 0.01 w /v% -CD ● : w ith 0.02 w /v% -CD ● : w ith 0.03 w /v% -CD ● : w ith 0.04 w /v% -CD ● : w ith 0.05 w /v% -CD ● : w ith 0.1 w /v% -CD ● : w ith 0.5 w /v% -CD ● : w ith 1 w /v% -CD ● : PBS
101
100 100
102
Shear stress (Pa)
(c)
● : HM-HPMC ● : w ith 0.01 w /v% -CD ● : w ith 0.02 w /v% -CD ● : w ith 0.03 w /v% -CD ● : w ith 0.04 w /v% -CD ● : w ith 0.05 w /v% -CD ● : w ith 0.1 w /v% -CD ● : w ith 0.5 w /v% -CD ● : w ith 1 w /v% -CD ● : PBS
101
102
Shear stress (Pa)
(d)
106
Strain (%)
104
102
100 100
● : HM-HPMC ● : w ith 0.01 w /v% -CD ● : w ith 0.02 w /v% -CD ● : w ith 0.03 w /v% -CD ● : w ith 0.04 w /v% -CD ● : w ith 0.05 w /v% -CD ● : w ith 0.1 w /v% -CD ● : w ith 0.5 w /v% -CD ● : w ith 1 w /v% -CD ● : PBS
101
102
Shear stress (Pa)
Fig. 6. Stress-strain curves of the 90L HM-HPMC hydrogel (0.5 w/v%) in the presence and absence of α-CD (a), β-CD (b), γ-CD (c) and changes in yield stress as a function of CD concentrations (d). Each point represents the mean ± S.E. of 3 experiments.
25
Fig. 7.
6h
0h
12 h
24 h
ICG alone
ICG +
HPMC ICG +
90L HM-HPMC +
-CD (0.04% w/v) Fig. 7.
In vivo fluorescence image after subcutaneous administration of an indocyanine green
(ICG) solution, the HPMC hydrogel (0.5 w/v%) containing ICG and the 90L HM-HPMC (0.5 w/v%)/-CD hydrogel containing ICG.
26
Fig. 8.
(b)
40
● : PBS ○ : insulin alone ● : HM-HPMC/ -CD hydrogel
30
＊
20 10 ＊ ＊
0 0
2
4
6
8
10
12
Plasma glucose level (mg/dl)
Plasma insulin level (ng/ml)
(a)
120 100 80 60 40
＊
● : PBS ○ : insulin alone ● : HM-HPMC/ -CD hydrogel
20 0 0
Time (h)
2
4
6
8
10
12
Time (h)
Fig. 8. Plasma levels for insulin (a) and glucose (b) after the subcutaneous administration of a phosphate buffered solution, an insulin solution (3 U/kg) and the 90L HM-HPMC (0.5 w/v%)/CD hydrogel (0.04 w/v%) containing insulin (3 U/kg) to rats. Each data point represents the mean ± S.E. of 3 experiments. *, p<0.05 versus insulin.
27
Credit Author Statement Masanori Okubo: Investigation, Writing- Original draft preparation, Validation, Formal Analysis. Daisuke Iohara: Investigation, Conceptualization, Writing- Reviewing and Editing, Supervision. Makoto Anraku: Resources. Taishi Higashi: Methodology. Kaneto Uekama: Visualization. Fumitoshi Hirayama: Writing- Reviewing and Editing, Project Administration.
Grafical abstract
28
Declaration of interests ☒ The authors declare that they have no known competing financial interests or personal relationships that could have appeared to influence the work reported in this paper.
☐The authors declare the following financial interests/personal relationships which may be considered as potential competing interests:
29
S0378-5173(19)30890-7 https://doi.org/10.1016/j.ijpharm.2019.118845 IJP 118845
To appear in:
International Journal of Pharmaceutics
Received Date: Revised Date: Accepted Date:
12 August 2019 18 October 2019 1 November 2019
Please cite this article as: M. Okubo, D. Iohara, M. Anraku, T. Higashi, K. Uekama, F. Hirayama, A Thermoresponsive Hydrophobically Modified Hydroxypropylmethyl Cellulose/Cyclodextrin Injectable Hydrogel for the Sustained Release of Drugs, International Journal of Pharmaceutics (2019), doi: https://doi.org/10.1016/ j.ijpharm.2019.118845
This is a PDF file of an article that has undergone enhancements after acceptance, such as the addition of a cover page and metadata, and formatting for readability, but it is not yet the definitive version of record. This version will undergo additional copyediting, typesetting and review before it is published in its final form, but we are providing this version to give early visibility of the article. Please note that, during the production process, errors may be discovered which could affect the content, and all legal disclaimers that apply to the journal pertain.
© 2019 Published by Elsevier B.V.
A Thermoresponsive Hydrophobically Modified Hydroxypropylmethyl
Cellulose/Cyclodextrin
Injectable Hydrogel for the Sustained Release of Drugs Masanori Okuboa, Daisuke Ioharaa,b*, Makoto Anrakua,b, Taishi Higashic, Kaneto Uekamaa, Fumitoshi Hirayamaa
aFaculty
of Pharmaceutical Sciences and bDDS Research Institute, Sojo University, 4-22-1
Ikeda, Nishi-ku, Kumamoto 860-0082, Japan cGraduate
School of Pharmaceutical Science, Kumamoto University, Oe-honmachi,
Kumamoto 862-0973, Japan
*Corresponding Author: Daisuke Iohara E-mail: [email protected] Tel: +81-96-326-5084 Fax: +81-96-326-5048
Abbreviations : Cyclodextrin (CD), Hydroxypropylmethyl cellulose (HPMC), Hydrophobically modified hydroxypropylmethyl cellulose (HM-HPMC), Degree of substitution (DS), Indocyanine green (ICG), Phosphate buffered saline (PBS), Maximum drug concentration (Cmax), Time required to reach maximum drug concentration (Tmax), Area under the concentration-time curve from time 0 to infinity (AUC0-∞), Mean residence time from time 0 to infinity (MRT0-∞), Storage modulus (G’), Loss modulus (G’’).
ABSTRACT The objective of this study was to develop a thermoresponsive injectable hydrogel for the sustained release of drugs by taking advantage of host-guest interactions between a hydrophobically modified hydroxypropylmethyl cellulose (HM-HPMC) and cyclodextrin (CD).
A thermoresponsive injectable hydrogel was prepared by simply adding CDs to HM-
HPMC hydrogel. The HM-HPMC hydrogel was converted into a sol with a low viscosity through host-guest interactions with CDs. The HM-HPMC/β-CD hydrogel became a gel near body temperature where the host dissociated from the hydrophobic moieties of the polymer in response to the temperature. The yield stress of the HM-HPMC became progressively lower on the addition of β-CD which was desirable in the case of developing an injectable formulation. When the HM-HPMC/-CD hydrogel containing indocyanine green (ICG) was subcutaneously administered to mice, the fluorescence of the ICG remained relatively constant for 24 h after the administration, which was substantially longer than that for ICG alone or an HPMC formulation. The plasma insulin level was maintained for a longer period of time when the HM-HPMC/-CD containing insulin was administered and the MRT value was increased by 1.6 times compared to a solution of insulin alone. In addition, the HM-HPMC/β-CD hydrogel formulation showed a prolonged hypoglycemic effect in response to the insulin which was slowly released from the hydrogel. A thermoresponsive injectable hydrogel was successfully constructed from the highly viscous HM-HPMC and β-CD, and the resulting formulation functioned as a sustained release carrier for drugs.
KEY WORDS cyclodextrin ･ thermoresponsive ･ injectable ･ hydrophobically modified polymer ･ hydrogel
2
1. Introduction In the past few years, numerous biotechnology-based medicines have been approved for treating various diseases (Morrison and Lahteenmaki, 2015). However, there are still many challenges in terms of developing the ideal formulation of a biotechnology-based medicine such as proteins because of their inherent characteristics related to their chemical and physical stability, immunogenicity and pharmacokinetics (Frokjaer and Otzen, 2005; Wu and Jin, 2008). Therefore, many biotechnology-based medicines except for antibody require daily or multiple injections for years or even an entire lifetime, which can lead to patient noncompliance. It is therefore very important to address such issues in the formulation design of the protein drugs as well as the low molecular weight drugs especially for the treatment of chronic diseases such as malignant tumors, diabetes, hepatitis, leukemia, rheumatoid arthritis etc. Hydrogels are composed of three dimensional networks of hydrophilic polymers in which the polymer chains are chemically or physically inter or intra molecularly cross-linked. Hydrogels have been frequently used as drug or cell delivery vehicles and have recently attracted considerable attention for use in tissue engineering because they possess high water contents and are able to mimic extracellular matrixes (Thambi et al., 2017; Nagahama et al., 2018; Lee and Mooney, 2001; Vashist et al., 2014; Gaharwar et al., 2014). In comparison with covalently linked hydrogels, physically linked hydrogels that can be prepared by inter or intra molecular interaction such as hydrophobic or related interactions, hydrogen bonding and host– guest interactions (Delplace et al., 2016), show sol-gel transitions in response to various external stimuli (pH, light, temperature and oxidizing/reducing agent etc.) (Sood et al., 2016; Harada et al., 2014; Qiu and Park, 2001; Yan et al., 2012). Among these stimuli-responsive hydrogels, thermoresponsive hydrogels are a classical example of a clinically approved formulation for use as a drug delivery vehicle (Takeuchi et al., 1999) because heating is an external stimulus which is readily available in every part of the body after administration. These 3
hydrogels are typically form sols at room temperature that have a low viscosity, but form gels when they are warmed to body temperature. Such injectable hydrogels are of great interest as potential materials for prolonging the retention of a drug at the administered site, leading to the modification of the release profile of the drug or absorption rate of the drug into body (Kempe and Maeder, 2012; Li et al., 2012; Yu and Ding, 2008; Alexander et al., 2013). Block copolymers such as poly(N-isopropylacrylamide) (Jeong et al., 2002; Soppimath et al., 2002; Alexander et al., 2014) or poly (ethylene oxide) (PEO)/poly(propylene oxide) (PPO) tri-block copolymers (Ruel-Gariepy and Leroux, 2004; Zhao et al., 2016; Park et al., 2009; Alexander et al., 2016) were, in the past, examined for this purpose, because such polymers undergo phase separation that is dependent on the environmental temperature (Parmar et al., 2017). However, higher polymer concentrations (above 20 w/v %) (Hoare and Kohane, 2008) was generally required for such a thermal gelation, which is costly and can cause allergic reactions. We recently developed a new strategy for preparing such thermoresponsive materials based on interactions between cyclodextrin (CD) and hydrophobically modified polymers (Iohara et al., 2017). Generally, the inclusion process for a CD is dependent on the temperature being used and is frequently occurs at low temperatures, while heating results in dissociation (Rekharsky and Inoue, 1998, 2000; Brocos et al., 2010). By simply adding CD to a hydrophobically modified hydroxypropylmethyl cellulose (HM-HPMC) hydrogel the gel became a low viscous sol as the result of the host-guest interaction of CD with the hydrophobic moiety. On heating, the CDs dissociate from the HM-HPMC/CD, resulting in the sol state turning into a gel state at around body temperature. On the basis of the previous study, we report here on the development of a thermoresponsive injectable hydrogel for the purpose of the sustained release of drugs. The HM-HPMC/CD is a sol with a low viscosity at low temperatures, which allows it to be easily injected, whereas it turns into a high viscous gel at the injection site (Fig. 1). An appropriate formulation of an injectable hydrogel was determined 4
by focusing on the thermoresponsive sol-gel transition property, and the release profile of insulin, a model drug, from the hydrogel was assessed after subcutaneous administration of the hydrogel.
2. Materials and method
2.1.
Materials
HM-HPMC (commercially known as Sangelose 60L or Sangelose 90L) and HPMC (METOLOSE 90SH - 30000) were obtained from the Daido Chemical Co. (Osaka, Japan). 60L HM-HPMC (M.W. 400,000) and 90L HM-HPMC (M.W. 600,000) contain a C-18 fatty acid group at the hydroxypropyl ends and has a degree of substitution (DS) of 0.3-0.6 mol %. The DS of the hydroxypropoxy and methoxy groups in 60L HM-HPMC were 7.0-11.0 mol % and 27.0-30.0 mol%, respectively. The 90L HM-HPMC contains hydroxypropoxy and methoxy groups 7.0-11.0 mol% and 21.5-24.0 mol%, respectively. The α-, - and -CD were obtained from Nihon Syokuhin Kako Co. (Tokyo, Japan). Indocyanine green (ICG) was purchased from Daiichi Sankyo Co. (Tokyo, Japan). Human insulin was purchased from Sigma Aldrich Co. (Tokyo, Japan). All other reagents were of analytical grade.
2.2.
Preparation of Thermoresponsive HM-HPMC/CD Injectable Hydrogels
HM-HPMC was added to a heated phosphate buffered saline (PBS) solution and dispersed by agitating. On cooling, the suspension of HM-HPMC was isolated as a clear HM-HPMC 5
hydrogel. Thermoresponsive HM-HPMC/CD hydrogels were prepared by simply adding various amount of α-, - and -CDs in the HM-HPMC hydrogels. The resulting HMHPMC/CD hydrogels were characterized by following rheological studies.
2.3.
Rheological Studies
Rheological investigations of the HM-HPMC/CD hydrogels were performed with a MCR-101 rheometer (Anton Paar Japan K.K., Tokyo, Japan). A cone and plate geometry with a diameter of 25 mm was used for these measurements. To evaluate the effect of CDs on the viscosity of HM-HPMC, steady shear rheology measurements were performed at 25.0 ºC with a shear rate at 0.1 s-1. Thermoresponsive property of HM-HPMC/CD hydrogels was evaluated by the viscosity measurement on the dynamic temperature and the dynamic rheology measurements. The viscosity of HM-HPMC/CD hydrogels was determined at a heating rate of 3.0 ºC/min at a shear rate of 1.0 s-1. The temperature of samples for the dynamic rheology was controlled to 10, 20, 30 and 37 ± 0.01 º C and the dynamic frequency data were obtained in the linear viscoelastic regime. For the usability test of the hydrogel for an injection, the measured strain was plotted as a function of shear stress.
2.4.
In Vivo Release Profile of ICG from the HM-HPMC/β-CD Hydrogel
A solution of the ICG was mixed with the HPMC or HM-HPMC/-CD hydrogel, and the final concentration of each component was fixed at 0.5 M for ICG, 0.5 w/v % for HPMC or HMHPMC and 0.04 w/v % for -CD, respectively. The HM-HPMC/-CD hydrogel containing 6
ICG was administered subcutaneously to the dorsal skin of ddy mice (male, 7 weeks of age) using a 1 ml syringe with a 26 gauge needle. The fluorescence intensity of ICG was measured using a fluorescence imaging apparatus (IVIS Lumina XR, PerkinElmeer Co., Waltham, MA, USA) at 0, 6, 12 and 24 h.
2.5.
Evaluation of HM-HPMC/β-CD Hydrogel for a Sustained Release of Insulin
A sample of the HM-HPMC (0.5 w/v %)/-CD (0.04 w/v %) product containing 3 U/kg of insulin was prepared for their administration. Insulin solution was simply mixed with HMHPMC/β-CD and the insulin did not affect the viscosity of the hydrogel. The hydrogel was administered subcutaneously to the dorsal skin of rats (male, 7 weeks of age) using a 1 ml syringe with a 26 gauge needle. Blood samples were collected from the tail vein at predetermined interval after the administration. Plasma insulin and glucose levels in the samples were measured using an Insulin ELISA KIT(TMB) (AKRIN-010T, Shibayagi, Gunma, Japan) or a Glucose-CII-Test Wako (Tokyo, Japan), respectively, following the manufacturer’s recommended protocol. A non-compartment model was used for the pharmacokinetic analysis. Pharmacokinetic parameters, including maximum drug concentration (Cmax), time required to reach maximum drug concentration (Tmax), area under the concentration-time curve from time 0 to infinity (AUC0- ∞ ) and mean residence time from time 0 to infinity (MRT0- ∞ ) were calculated by the moment analysis program which is available on Microsoft Excel.
2.6.
Statistical Analysis
7
Data are presented as the mean values from n samples, and the results are reported as the mean ± S.E. Kruskal-Wallis analysis of variance (ANOVA) was performed followed by student’s ttests for comparison of samples. For all analyses, values of p < 0.05 were regarded as statistically significant.
3. Result and discussion
3.1. Preparation and Evaluation of Thermoresponsive HM-HPMC/CD Injectable Hydrogel for Injection
3.1.1. Modification of the Viscosity of HM-HPMC Hydrogels by CDs HM-HPMC samples with different molecular weights (60L HM-HPMC and 90L HMHPMC) were tested as the base of the injectable hydrogel. HM-HPMC solutions are highly viscous due to the intra- and inter-molecular associations of the stearyl groups which are introduced at the hydroxypropyl ends (Ghosal et al., 2012; Ikeda et al., 1994). We previously reported that the high viscosity of a 60L HM-HPMC hydrogel was markedly decreased in the presence of small amounts of α-CD (Iohara et al., 2017). The small cavity of the α-CD allowed the stearyl moiety of the HM-HPMC to be selectively incorporated, which inhibited the associations of the polymers. The viscosity of the 60L and 90L HM-HPMC hydrogels was first compared in the presence of CDs in order to design an appropriate formulation for an injectable hydrogel. Figure 2 shows the viscosity of the 60L and 90L HM-HPMC hydrogels at a shear rate of 0.1 s-1. The high viscosity was observed for both the 60L and 90L HM-HPMC hydrogels alone at low concentration (0.5 w/v %), the values of which were about 34 Pa･s for 60L HMHPMC and 53 Pa･s for 90L HM-HPMC. Such a high viscosity of the 60L and 90L HM-HPMC 8
decreased by the addition α-CD or β-CD, and the reducing effect was dependent on the added amount of CDs. In the case of -CD, the reducing effect on the viscosity was small and a high viscosity was observed, even when 0.2 w/v% of -CD was added. In comparing the viscosity among the two HM-HPMC samples, the viscosity of the 90L HM-HPMC was higher than that of the 60L HM-HPMC over the entire range of CD concentrations. This can be attributed to the fact that the 90L HM-HPMC with a higher molecular weight contains a higher number of stearyl moieties (15 molecules per polymer chain) compared with 60L HM-HPMC (10 molecules per polymer chain). The viscosity of 90L HM-HPMC as well as 60L HM-HPMC was modified by the addition of α- and β-CD. A hydrogel with a higher viscosity would be expected to release drugs more slowly from its gel networks and for this reason, we chose to use the 90L HM-HPMC as the base for the injectable hydrogel.
3.1.2. Thermoresponsive Changes in Viscosity of the 90L HM-HPMC/CD Hydrogel Figure 3 shows the change in viscosity of the 90L HM-HPMC as a function of dynamic temperature. The high viscosity of the 90L HM-HPMC alone decreased with increasing temperature in the range of 20 to 45ºC similar to the behavior of a typical elastic polymer solution. In contrast, the viscosity of the 90L HM-HPMC/α-CD was low at low temperatures, but the viscosity increased slightly with temperature in the range of 40~45ºC in the presence of 0.04% α-CD. Such changes in viscosity were minor in the presence of 0.1% or 0.2% α-CD. Regarding the -CD, the viscosity clearly changed with increasing temperature, reaching a maximum viscosity at around 43ºC at 0.04 w/v% -CD. In the presence of -CD, a similar viscosity change was observed for 0.2 w/v% -CD and the maximum viscosity occurred at approximately 36ºC. We previously reported that the interaction of CD with a C-18 fatty acid is dependent on the size of the CD cavity; thus, the α-CD with a small cavity size showed a high affinity for the stearyl moiety, followed by the -CD and the -CD. In addition, the 9
interaction between the CD and the stearyl moiety varied with temperature and the dissociation appeared to occur at a higher temperature (Fig. S1). The different rheological behavior of the 90L HM-HPMC in the presence of the CDs can be explained by the fact that the viscosity of such systems is controlled by the host-guest interactions. The viscosity of the 90L HMHPMC/α-CD barely changed on heating because the stearyl moieties were contained within the α-CD which has a relatively higher affinity for stearyl moieties, even at higher temperatures. In the case of β-CD, a clear viscosity change was observed at 0.04% β-CD, where the dissociation of the β-CD occurred on heating. Because -CD has a low affinity for stearyl moieties, they readily dissociate, leading to a maximum viscosity at lower temperatures compared with the β-CD system. These rheological behaviors were different from the previously reported 60L HM-HPMC/CDs system (Fig. S1). The peak for the maximum was shifted to a higher temperature in the case of 90L HM-HPMC, probably because of the large number of stearyl moieties per molecule. The viscosity of the 90L HM-HPMC/-CD hydrogel was measured upon heating and cooling cycles (Fig. 4). The 90L HM-HPMC (0.5 w/v%)/βCD (0.04 w/v%) hydrogel was used for the measurement because it showed a maximum viscosity near body temperature. The viscosity of the 90L HM-HPMC/-CD hydrogel was low below 1.0 Pa･s at 10ºC, whereas it was six times as high at 37ºC. The changes in viscosity were markedly larger compared with those of the previous 60L HPMC/α-CD system (2.0 Pa‧ s at low temperature versus 4.0 Pa‧s at high temperature). The change was reversible upon heating and cooling because the inclusion process of the CD is in an equilibrium state. Temperature dependent viscosity changes were evaluated by dynamic rheology measurements (Fig. 5). The 90L HM-HPMC/-CD hydrogel showed a tan (loss modulus G’’/storage modulus G’) of > 1 at 10ºC, indicating that it was in a sol state. It should be noted that the tan value was ≒ 1 at 20ºC, indicating that it was in a sol-gel intermediate state. The tan values became < 1 at 30ºC and 37ºC, indicating that they were in a gel state. These collective results 10
suggest that the 90L HM-HPMC/-CD hydrogel was converted from a sol into a gel in response to a temperature change. A thermoresponsive sol-gel transition system could be constructed with 90L HM-HPMC by adding a CD with the appropriate cavity size and concentration.
3.1.3.
Rheological Properties of 90L HM-HPMC/CDs Hydrogel for an Injection
A hydrogel is generally a useful carrier for the sustained release of drugs, but the high viscosity of the hydrogel limits its use in practical applications. For use of a hydrogel in the formulation of an injectable formulation, the hydrogel must be readily ejected from the needle of the syringe. The rheological properties of the 90L HM-HPMC/CD hydrogels were evaluated in terms of the usability of the preparation for an injection. The strain of samples was plotted against shear stress, as shown in Fig. 6. Regarding the 90L HM-HPMC alone, the strain rapidly increased at 16 Pa indicating that a reasonable amount of force was required to collapse the gel structure and to initiate a flow. Such a sharp change in strain could be eliminated when the above system was used and a pattern similar to a PBS could be achieved. The value for shear stress at the inflection point (yield stress) was plotted against CDs concentrations in Fig 6(d). The yield stress of the 90L HM-HPMC hydrogel was decreased by adding CDs, and α-CD and β-CD showed a remarkable decrease in yield stress at low concentrations, i.e., when 0.02 w/v% α-CD or 0.05 w/v% -CD were added. Regarding the -CD, the effect on yield stress of the 90L HM-HPMC was weak and at least 0.5 w/v% was needed to decrease the yield stress. These results indicate that a 90L HM-HPMC/CD hydrogel with a low yield stress would be applicable to the injectable formulation. The viscosity of the 90L HM-HPMC/CDs hydrogels was also measured as a function of share rate (Fig. S2). The viscosity decreased with share rate showing quasi viscous flow properties. As expected, α-CD and -CD reduced the viscosity of 90L HMHPMC in the concentration dependent manner. On the basis of these collective data, we used 11
the 90L HM-HPMC (0.5 w/v%)/β-CD (0.04 w/v%) hydrogel for the formulation of the injectable hydrogel because the system showed a low yield stress and viscosity at low temperature as well as undergoing thermal gelation near body temperature (Fig. 3). α-CD was useful to modify the viscosity of the 90L HM-HPMC, however, the system did not undergo thermal gelation near body temperature. In the case of the γ-CD system, high CD concentrations were needed to modify the viscosity of the 90L HM-HPMC. Although the HMHPMC (0.5 w/v%)/γ-CD (0.2 w/v%) hydrogel underwent thermal gelation near body temperature, the yield stress and the viscosity of the formulation (Fig. 6 and Fig. S2) were not sufficiently adaptable for the formulation of the injectable hydrogel compared to the β-CD formulation. The 90L HM-HPMC (0.5 w/v%)/β-CD (0.04 w/v%) hydrogel could be readily ejected from the syringe with a 27G needle, whereas this was not possible for the HM-HPMC alone (Movie S1). The HM-HPMC (0.5 w/v%)/β-CD (0.04 w/v%) hydrogel, which undergoes thermal gelation near body temperature, was used as the base for the injectable hydrogel in the following in vivo studies.
3.2.
In Vivo Profile for the release of a Fluorescent Dye from the 90L HM-HPMC/β-CD
Hydrogel The release properties of the drug from the 90L HM-HPMC (0.5 % w/v)/-CD (0.04 % w/v) hydrogel was evaluated in in vivo studies. The 90L HM-HPMC/-CD hydrogel containing indocyanine green (ICG) was administered subcutaneously to mice and the fluorescence of the ICG was monitored at predetermined intervals (Fig. 7). The fluorescence of the ICG decreased dramatically within 12 h after administration and disappeared by 24 h in the case of the ICG solution alone. This result indicate that ICG rapidly diffused into the body from the solution. Similar behavior was observed for the HPMC formulation. On the other hand, a strong 12
fluorescence continued to be observed at 12 h after the administration of the 90L HM-HPMC/CD hydrogel, and fluorescence was also observed even in 24 h. Thermal gelation occurred as soon as the hydrogel was warmed to body temperature. Thus, it was estimated that the thermal gelation occurred within a few minutes after its administration because the injection volume (0.1 ml) was small enough to permit it to be heated rapidly. It should be noted here that the viscosity of the 90L HM-HPMC/β-CD hydrogel was essentially the same as that of HPMC at room temperature (Fig. S3). It can thus be concluded that the extended release of ICG from the 90L HM-HPMC/β-CD hydrogel can be attributed to the thermosresponsive sol-gel transition on being warmed to body temperature.
3.3. Evaluation of the Thermoresponsive 90L HM-HPMC/-CD Hydrogel for the Sustained Release of a Drug In order to evaluate the sustained-release property of the 90L HM-HPMC/-CD hydrogel, insulin (as a model drug) was loaded into the hydrogel. We confirmed that insulin had no effect on the rheological properties of the hydrogel (data not shown). Figure 8(a) shows the plasma insulin concentrations following the subcutaneous administration of 90L HM-HPMC/-CD hydrogel containing insulin. The maximum levels for plasma insulin were reached within 1 h after the administration of an insulin in an ordinary solution (Cmax: 42.8 ± 2.8 ng/ml), and the administered insulin disappeared from the blood circulation within 4 h. When the insulin in the 90L HM-HPMC/β-CD hydrogel formulation was administered, the insulin concentration also reached the maximum level within 1 h after administration (Cmax: 33.6 ± 1.7 ng/ml), but, in this case, the plasma insulin level at 4 h and 6 h after the administration was significantly higher than that of the insulin alone solution. The pharmacokinetics data for these formulations are shown in Table 1. The MRT value for the 90L HM-HPMC/β-CD hydrogel was prolonged by 13
about 1.6 times compared with the insulin alone solution (90L HM-HPMC/-CD = 2.6 ± 0.2 h, insulin alone solution = 1.6 ± 0.1 h). The plasma glucose levels were also monitored after the administration of 90L HM-HPMC/-CD hydrogel (Fig. 8(b)). When the insulin alone solution was administered to rats, the plasma glucose level dropped to about 40 mg/dl at 2 h after the administration, and the value recovered to normal levels within 6 h. In sharp contrast, the low plasma glucose level was maintained even at 6 h in the case of the administration of the 90L HM-HPMC/-CD hydrogel formulation (40 mg/dl at 6 h). The hypoglycemic effect was obviously prolonged for the 90L HM-HPMC/-CD hydrogel formulation. It was difficult to administer the insulin in the HM-HPMC formulation as shown in Movie S1. The HM-HPMC was converted into a sol with a low viscosity when the β-CD was added that permitted it to be administered from the syringe. The administered 90L HM-HPMC/β-CD hydrogel became a highly viscous gel by warming at body temperature, thus the insulin was slowly released from the gel matrix resulting in a prolonged hypoglycemic effect.
Table 1. In vivo pharmacokinetic parameters of insulin and the 90L HM-HPMC (0.5 w/v%)/CD (0.04 w/v%) hydrogel containing insulina AUC0-∞d
Sample
Cmaxb (ng/ml)
Tmaxc (h)
Insulin alone
42.8 ± 2.8
0.8 ± 0.2
79.0 ± 6.2
1.6 ± 0.1
33.6 ± 1.7*
0.5 ± 0.0
85.2 ± 3.5
2.6 ± 0.2*
Insulin/ 90L HM-HPMC/-CD aEach
(ng/ml･h)
value represents the mean ± S.E. of 3 experiments.
bMaximum
drug concentration.
cTime
required to reach maximum drug concentration.
dArea
under the curve from time 0 to infinity.
eMean
residence time from time 0 to infinity.
*p < 0.05 versus insulin. 14
MRT0-∞e (h)
4. Conclusions Thermoresponsive injectable hydrogels were successfully prepared by using a combination of the highly viscous 90L HM-HPMC and β-CD. The 90L HM-HPMC/β-CD underwent thermal gelation near body temperature as the result of host–guest interactions between the stearyl moieties of HM-HPMC and β-CD cavity. The 90L HM-HPMC/β-CD hydrogel was easy to administer by simple injection. The 90L HM-HPMC/β-CD turned into a high viscous gel after being entering the body, which allowed the drugs to be slowly released and efficacy to be maintained. The thermal response of the HM-HPMC/CD material can be tuned by selecting an appropriate CD and the concentrations of the components. A thermoresponsive injectable hydrogel, in which interactions between the CD and hydrophobically modified polymers determine its properties, promises to be useful as an advanced drug carrier for drugs incluiding proteins
Declaration of competing interest None.
Figure captions
Fig. 1.
Formulation of the sustained release system using a thermoresponsive HM-
HPMC/CD hydrogel.
Fig. 2. Changes in the viscosity of 60L and 90L HM-HPMC hydrogels (0.5 w/v%) by adding CDs. The shear viscosity at 0.1 s-1 was plotted as a function of CDs concentrations. Each point represents the mean ± S.E. of 3 experiments. 15
Fig. 3. Changes in the viscosity of the 90L HM-HPMC/CDs hydrogel (0.5 w/v%) on the dynamic temperature. The viscosity was determined at a shear rate of 1.0 s-1. Each point represents the mean ± S.E. of 3 experiments. Fig. 4. Thermoresponsive reversible viscosity change of the 90L HM-HPMC (0.5 w/v%)/CD (0.04 w/v%) hydrogel. The viscosity was determined at a shear rate of 1.0 s-1. Each point represents the mean ± S.E of 3 experiments.
Fig. 5. Loss tangents (tan ) for the 90L HM-HPMC (0.5 w/v%)/-CD (0.04 w/v%) hydrogel at 10, 20, 30 and 37℃. Each point represents the mean ± S.E of 3 experiments. ※tan (loss tangent) = G’’ (loss modulus) / G’ (storage modulus)
Fig. 6. Stress-strain curves of the 90L HM-HPMC hydrogel (0.5 w/v%) in the presence and absence of α-CD (a), β-CD (b), γ-CD (c) and changes in yield stress as a function of CD concentrations (d). Each point represents the mean ± S.E. of 3 experiments. Fig. 7. In vivo fluorescence image after subcutaneous administration of an indocyanine green (ICG) solution, the HPMC hydrogel (0.5 w/v%) containing ICG and the 90L HM-HPMC (0.5 w/v%)/-CD hydrogel containing ICG.
Fig. 8. Plasma levels for insulin (a) and glucose (b) after the subcutaneous administration of a phosphate buffered solution, an insulin solution (3 U/kg) and the 90L HM-HPMC (0.5 w/v%)/-CD hydrogel (0.04 w/v%) containing insulin (3 U/kg) to rats. Each data point represents the mean ± S.E. of 3 experiments. *, p<0.05 versus insulin.
Acknowledgement
This work was supported, in part, by JSPS KAKENHI Grant Number 18K06618. The authors are grateful to Mr. Yamada, Ms. Take and Ms. Miyata for technical assistance.
16
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19
Fig. 1.
HM-HPMC R = -H -CH3 -CH2CH(OH)CH3
addition of CDs
HM-HPMC/CD
-CH2CH(OH)CH2OC18H37
heating OH
heating
O
HO
OH
O OH
O
O HO
HO
OH O HO
OH
O HO
O OH
OH OH
O OH
cooling
addition of CDs
OH
O OH
O
OH
heating
HO
heating
gel
sol
O
O
α-CD (n = 1) β-CD (n = 2) γ-CD (n = 3)
Subcutaneous injection
sol
Body temperature cooling (37℃)
gel
Easy to inject Sustained release of drug Fig. 1. Formulation of the sustained release system using a thermoresponsive HM-HPMC/CD hydrogel.
20
Fig. 2.
60
60
50
○：60L HM-HPMC ●：90L HM-HPMC
40
Viscosity (Pa・s)
Viscosity (Pa・s)
50
30 20
○：60L HM-HPMC ●：90L HM-HPMC
40 30 20 10
10 0
0 0
0.05
0.1
0.15
0.2
0
0.05
Concn. of α-CD (w/v%)
0.1
0.15
Concn. of β-CD (w/v%)
60
○：60L HM-HPMC ●：90L HM-HPMC
Viscosity (Pa・s)
50 40 30 20 10 0 0
0.05
0.1
0.15
0.2
Concn. of γ-CD (w/v%)
Fig. 2. Changes in the viscosity of 60L and 90L HM-HPMC hydrogels (0.5 w/v%) by adding CDs. The shear viscosity at 0.1 s-1 was plotted as a function of CDs concentrations. Each point represents the mean ± S.E. of 3 experiments.
21
0.2
Fig. 3.
Fig. 3. Changes in viscosity of the 90L HM-HPMC/CDs hydrogel (0.5 w/v%) on the dynamic temperature. The viscosity was determined at a shear rate of 1.0 s-1. Each point represents the mean ± S.E. of 3 experiments.
22
Fig. 4.
7
Viscosity (Pa･s)
6 5 4 3 2 1 0 10
37
10
37
10
37
10
37
10
Temperature (℃)
Fig. 4. Thermoresponsive reversible viscosity change of the 90L HM-HPMC (0.5 w/v%)/-CD (0.04 w/v%) hydrogel. The viscosity was determined at a shear rate of 1.0 s-1. Each point represents the mean ± S.E of 3 experiments.
23
Fig. 5.
Fig. 5. Loss tangents (tan ) for the 90L HM-HPMC (0.5 w/v%)/-CD (0.04 w/v%) hydrogel at 10, 20, 30 and 37℃. Each point represents the mean ± S.E of 3 experiments. tan (loss tangent) = G’’ (loss modulus) / G’ (storage modulus)
24
Fig. 6.
(b)
106
106
104
104
102
100 100
Strain (%)
Strain (%)
(a)
102
● : HM-HPMC ● : w ith 0.01 w /v% -CD ● : w ith 0.02 w /v% -CD ● : w ith 0.03 w /v% -CD ● : w ith 0.04 w /v% -CD ● : w ith 0.05 w /v% -CD ● : w ith 0.1 w /v% -CD ● : w ith 0.5 w /v% -CD ● : w ith 1 w /v% -CD ● : PBS
101
100 100
102
Shear stress (Pa)
(c)
● : HM-HPMC ● : w ith 0.01 w /v% -CD ● : w ith 0.02 w /v% -CD ● : w ith 0.03 w /v% -CD ● : w ith 0.04 w /v% -CD ● : w ith 0.05 w /v% -CD ● : w ith 0.1 w /v% -CD ● : w ith 0.5 w /v% -CD ● : w ith 1 w /v% -CD ● : PBS
101
102
Shear stress (Pa)
(d)
106
Strain (%)
104
102
100 100
● : HM-HPMC ● : w ith 0.01 w /v% -CD ● : w ith 0.02 w /v% -CD ● : w ith 0.03 w /v% -CD ● : w ith 0.04 w /v% -CD ● : w ith 0.05 w /v% -CD ● : w ith 0.1 w /v% -CD ● : w ith 0.5 w /v% -CD ● : w ith 1 w /v% -CD ● : PBS
101
102
Shear stress (Pa)
Fig. 6. Stress-strain curves of the 90L HM-HPMC hydrogel (0.5 w/v%) in the presence and absence of α-CD (a), β-CD (b), γ-CD (c) and changes in yield stress as a function of CD concentrations (d). Each point represents the mean ± S.E. of 3 experiments.
25
Fig. 7.
6h
0h
12 h
24 h
ICG alone
ICG +
HPMC ICG +
90L HM-HPMC +
-CD (0.04% w/v) Fig. 7.
In vivo fluorescence image after subcutaneous administration of an indocyanine green
(ICG) solution, the HPMC hydrogel (0.5 w/v%) containing ICG and the 90L HM-HPMC (0.5 w/v%)/-CD hydrogel containing ICG.
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Fig. 8.
(b)
40
● : PBS ○ : insulin alone ● : HM-HPMC/ -CD hydrogel
30
＊
20 10 ＊ ＊
0 0
2
4
6
8
10
12
Plasma glucose level (mg/dl)
Plasma insulin level (ng/ml)
(a)
120 100 80 60 40
＊
● : PBS ○ : insulin alone ● : HM-HPMC/ -CD hydrogel
20 0 0
Time (h)
2
4
6
8
10
12
Time (h)
Fig. 8. Plasma levels for insulin (a) and glucose (b) after the subcutaneous administration of a phosphate buffered solution, an insulin solution (3 U/kg) and the 90L HM-HPMC (0.5 w/v%)/CD hydrogel (0.04 w/v%) containing insulin (3 U/kg) to rats. Each data point represents the mean ± S.E. of 3 experiments. *, p<0.05 versus insulin.
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Credit Author Statement Masanori Okubo: Investigation, Writing- Original draft preparation, Validation, Formal Analysis. Daisuke Iohara: Investigation, Conceptualization, Writing- Reviewing and Editing, Supervision. Makoto Anraku: Resources. Taishi Higashi: Methodology. Kaneto Uekama: Visualization. Fumitoshi Hirayama: Writing- Reviewing and Editing, Project Administration.
Grafical abstract
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Declaration of interests ☒ The authors declare that they have no known competing financial interests or personal relationships that could have appeared to influence the work reported in this paper.
☐The authors declare the following financial interests/personal relationships which may be considered as potential competing interests:
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