1 Introduction

Malignant tumor is a major hazard to human life and health safety because of its intricacy. A number of therapeutic approaches for malignancies have emerged in the clinic to address this issue, including conventional chemoradiotherapy [1], photothermal therapy [2] immunotherapy [3] and chemokinetic therapy [4], among others. However, a single therapeutic method gradually becomes unable to meet the goal of maximizing the effect of antitumor. The proposed combination therapy for tumors can maximize the anti-tumor effect, and the development of nanotechnology offers a practical platform for combination therapy, which can increase the tumor uptake of chemotherapeutic drugs and reduce off-target toxicity [5], thereby obtaining a better therapeutic effect. Among the many combination therapies, the combination of photothermal therapy and chemotherapy has gained popularity as a kind of treatment. Based on chemotherapy, photothermal therapy [6] uses heat to destroy tumor cells by converting the light energy absorbed by the photothermal agent into heat. The heat increases the fluidity of the cell membrane and improves the cells’ ability to absorb chemotherapeutic medications. Currently, researchers are committed to developing various photothermal agents [7, 8] to realize the conjunction of chemotherapy and photothermal therapy, such as inorganic photothermal agents (noble metals, carbon-based materials, etc.) possessed high stability and strong photothermal conversion efficiency, and organic photothermal agents (indocyanine green, organic polymer nanoparticles, etc.) with good biosafety and degradation. Bao et al. [9] carried doxorubicin hydrochloride (DOX) by coating the surface of bismuth nanoparticles with silica (BMSN) and found that the cell survival rate of the photothermal combination chemotherapy group after the introduction of NIR light was significantly lower than that of alone, and the mice’s tumors in the combination group were significantly regressed, which showed a significant anti-tumor effect. Li et al. [10] created a new type of tumor-targeting cell membrane capsule to carry doxorubicin hydrochloride and indocyanine green for combined chemotherapy with photothermal therapy to treat cancer. They found that the structure could notably hinder the growth of tumors in xenografts in mice in the existence of near-infrared light without significant toxicity.

Even though chemotherapy is still an essential part of the treatment of tumors nowadays, the majority of chemotherapeutic drugs have non-specific distribution in the organism that can cause toxic side effects in healthy tissues and are not suitable for internal transport because of their inherent properties [11] The photothermal agents used in photothermal therapy have the characteristics of single function and poor chemical stability [12]. The application of nanotechnology has largely solved this problem by improving the delivery and targeting of chemotherapeutic drugs and photothermal agents through the design of rational nanoplatforms to enhance the anti-tumor effect. The outstanding catalytic properties and stable physicochemical features of titanium dioxide make it widely used in the domains of sensing, catalytic degradation, UV reflection, and drug delivery [13]. To achieve the conjunction of photothermal and chemotherapy, titanium dioxide, which has high photocatalytic characteristics, is employed as a substrate material in photothermal therapy. Behnam et al. [14] synthesized pegylated titanium dioxide nanoparticles and evaluated their inhibitory effect on tumors under laser excitation therapy. The outcomes demonstrated that the nanoparticles used in photothermal therapy were particularly efficient at eliminating solid tumors. A low-cost and high-performing photothermal agent, Ag@TiO2, was created by Nie et al. [15] which had an outstanding near-infrared (NIR) absorption of 808 nm and an excellent photothermal conversion efficiency of 65%. They tested the photothermal cytotoxicity of the particles in mice, and the outcomes showed that irradiation with a near-infrared laser (2 W/cm2) for only 1 min led to a decrease in the tumor volume and a significant anti-tumor effect. Based on the cross-application of nanotechnology and biomedicine, increasing the amount of chemotherapeutic drugs carried by nanocarriers can effectively enhance the anti-tumor effect, so it is necessary to carry more chemotherapeutic drugs by constructing hollow mesoporous titanium dioxide nanoparticles.

Studies have shown that titanium dioxide may have associated biotoxicity risks. Liu et al. [16] discovered that mice given large doses of titanium dioxide experienced severe harm to their liver, kidney, and myocardium in addition to abnormalities in their blood glucose and lipid balance, and triggered related inflammatory responses. Sun et al. [17] discovered that prolonged contact with titanium dioxide nanoparticles caused lung inflammation and hemorrhage, which was connected to the fact that the particles encouraged the expression of inflammatory cytokines in the lungs and induced the overindulgent generation of reactive oxygen species in the lungs. However, the damage caused by titanium dioxide nanoparticles to the organism is essentially mediated by a large number of nanoparticles [16, 18]. E. Fabian et al. [19] exposed rats to lower doses of TiO2 nanoparticles to explore their tissue distribution and toxicity, and found that a dose of 5 mg/kg TiO2 provided by intravenous administration had a 100% bioavailability, with no significant toxic effects in experimental animals, which revealed that that TiO2 nanoparticles can be used safely at low doses. And careful surface modification can modulate the biotoxicological behavior of TiO2, thereby eliminating or reducing potential inflammatory responses after administration.

Mesoporous silica (MSN) is gradually used in many important applications due to its unique structural properties such as high specific surface area, ordered mesoporous structure, wide pore size distribution (2–50 nm), and nanoscale size. Research fields such as adsorption, catalysis, and biomedicine have shown broad application prospects, and their large-scale synthesis strategies [20, 21] are constantly being optimized due to their wide application. However, some studies have shown that mesoporous silica can induce varying degrees of toxicity. Z. Tao et al. [22] studied the impact of mesoporous silica nanoparticles on cell bioenergy and found that MSN can inhibit cellular and mitochondrial respiration. W. Lin et al. [23] investigated the cytotoxicity of silica nanoparticles in co-incubation with human bronchioloalveolar carcinoma-derived cells and found that cell viability was reduced dose-dependently. Quantitative assessment of relevant indicators revealed that this was associated with increased oxidative stress. Studies [24, 25] have shown that bare mesoporous silica can cause erythrocyte lysis and further cause hemolysis. This is because MSN can electrostatically interact with the positively charged erythrocyte membrane through its negatively charged surface silanol groups. Both toxicity and hemolysis are closely related to the surface of nanoparticles, so the toxicity and hemolytic activity of MSN can be improved by modifying its surface. A. Yildirim et al. [26] found that the modified MSN can reduce or eliminate the hemolytic activity of MSN by modifying the surface of MSN with organosilanes, rhodamine, and polyethylene glycol. The introduction of specific organic groups improved the biocompatibility of mesoporous silica to some extent, such as reducing the hemolytic activity, cytotoxicity and improving the biodegradability of the material. Periodic mesoporous organosilica [27] (PMO) possessing organic groups has the advantages of excellent biocompatibility, easy surface functional modification and control of drug release, and is used in various fields of biomedicine. Coating the surface of titanium dioxide with PMO can greatly improve the biocompatibility and reduce the biological toxicity of titanium dioxide.

In order to maximize the anti-tumor effect, the nanoplatform needs to be targeted so that it can accumulate in large amounts within the tumor tissue to exert the optimal anti-tumor effect. Hyaluronic acid (HA) is a linear anionic polymer [28]. Its structural properties determine that it can easily combine with other chemicals. HA possesses excellent biocompatibility, biodegradability, and receptor-binding properties. HA [29] can specifically bind to the receptors of CD44 and hyaluronate mediated motility (RHAMM) which are overindulgently expressed on the surface of a variety of tumor cells and are able to bind to the nanoparticles modified with HA to achieve the targeted delivery of antitumor drugs through the ligand-receptor binding mechanism. In addition, the functional modification of HA can extend the blood circulation time [30, 31] and increase the accumulation of tumors to a certain extent, which avoids the accumulation of more nanoparticles in organs and reduces body toxicity.

In this paper, hollow mesoporous titanium dioxide was prepared and coated with PMO to carry the chemotherapy drug DOX and the photothermal agent ICG, and then functionalized with HA to achieve targeted delivery and prevent drug leakage during transportation (Scheme 1). The system was a pH/NIR dual-response chemo-photothermal platform. After the HMTNs@PMO-DOX/ICG@HA targeted into breast cancer cells (MCF-7), the release of photothermal agents and chemotherapeutic drugs was achieved by the degradation of HA in a slightly acidic environment, and the introduction of near-infrared (NIR) light caused the tumor temperature to rise locally and resulted in thermal damage, which in turn exerted a direct killing effect of the chemotherapeutic drug DOX. The above measures achieved combined chemotherapy-photothermal therapy and thus enhanced the anti-tumor effect.

Scheme 1
scheme 1

Preparation principle and synergistic treatment process of HMTNs@PMO-DOX/ICG@HA

Full size image

2 Experimental methods

2.1 Materials

Tetraethyl orthosilicate (TEOS, AR), sodium hydroxide (NaOH, AR) and ammonium hydroxide (25–28%) were purchased from China Chengdu. Ethylalcohol (EtOH, AR, China Tianjin). Hyaluronic acid (HA, 40-100KDa), titanium butoxide (TBOT, AR), bis (triethoxysilyl) ethane (BTEE, AR), 3-aminopropyltriethoxysilane (APTES, AR), N-(3-dimethylaminopropyl) -N’-ethylcarbodiimide hydrochloride (EDC, 98.5%) and N-hydroxysuccinimide (NHS, 99%) were purchased from Macklin. Hydroxypropyl cellulose (HPC, M.W:100000), phosphate buffer saline (PBS, pH = 7.4 and pH = 5.0), doxorubicin (DOX) and indocyanine green (ICG) were purchased from Aladdin. Cell Counting Kit-8 (CCK-8) was purchased from Invigentech. Cetyltrimethylammonium bromide (CTAB, 99%) and polyvinylpyrrolidone (PVP, M.W: 40,000) were purchased from Solarbio. All chemicals were used without further purification. All of the trials utilized deionized water.

2.2 Characterization

The morphology and size of the samples were observed by scanning electron microscopy (SEM, JSM7600F, JEOL, Nippon Electron Co., Ltd.) and transmission electron microscope (TEM, Zeiss FEIG20, FEI, Inc., USA). The crystal structure and phase of the materials were analyzed by X-ray diffraction (XRD, D8 Adcance, Brooke, Germany). Nitrogen adsorption-desorption isotherms and pore size distributions were determined using a specific surface and porosity analyzer (BET, Micromeritics, Mack Instruments, USA). Zeta potential of the nanoparticles were characterized using a NanoBrook (Brookhaven Instruments Inc., USA). The absorbance of the drugs in the samples was determined by an ultraviolet-visible spectrophotometer (UV–Vis, UV-6100s, Shanghai Mepda Instrument Co., Ltd.).

2.3 Synthesis of HMTNs

Silica-protected calcination and alkali etching were used to create hollow mesoporous titanium dioxide nanoparticles (HMTNs) [32]. According to a previously reported procedure [33], a silica core was prepared as a template. Next, titanium dioxide was wrapped on the silica core’s surface using the sol-gel method, and then a layer of silica was added as a protective layer. Finally, HMTNs were produced by calcination and alkali etching.

2.4 Synthesis of HMTNs@PMO-NH2

0.085 g of HMTNs were ultrasonically dispersed in a mixed solution containing 5.0 mL of anhydrous ethanol, 75.0 mL of deionized water, 0.150 g of CTAB, and 1.8 mL of ammonia, and the reaction was subsequently carried out at 30 °C for 1 h. The reaction was continued for 3 h after adding 80 µL of BTEE drop by drop and the products were centrifuged, washed and dried. The HMTNs@PMO were obtained by removing the template CTAB through refluxing the product for 24 h at 60 °C using 10.0 mL anhydrous ethanol and 1.0 mL concentrated hydrochloric acid [34].

0.1 g HMTNs@PMO were ultrasonically dispersed in 60.0 mL of anhydrous ethanol, and 20.0 mL of anhydrous ethanol solution with 2.0 mL of APTES was added, and the reaction was carried out at 25 °C for 24 h. HMTNs@PMO-NH2 was acquired by centrifuging, washing and drying.

2.5 Synthesis of HMTNs@PMO-DOX/ICG@HA

0.036 g HMTNs@PMO-NH2 were ultrasonically dispersed in 90 mL of DOX aqueous solution (0.4 mg/mL) added with 6.0 mL PBS buffer (pH = 7.4), and they were ultrasonicated for 20 min and stirred at 25 °C for 48 h in the dark. When the reaction time was up, the mixed solution was centrifuged and washed with deionized water until the supernatant was clear. All supernatants’ absorbance values at 480 nm were calculated. The HMTNs@PMO-DOX finally got after drying under vacuum at 40 °C for 24 h, and the drug loading rate was determined according to formula (1) [35].

$${\text{wt}}\% {\text{ }} = {\text{ }}\left[ {\frac{{m_{1} - CV}}{{m_{2} {\text{ }} + {\text{ }}\left( {m_{1} - CV} \right)}}} \right] \times 100\%$$
(1)

In the formula, m1: the amount of the initial drug, mg; m2: the amount of the carrier material dispersed in the suspension, mg; C: the concentration of the drug in the supernatant, mg/mL; V: the volume of the suspension, mL.

0.020 g HMTNs@PMO-DOX were dispersed in 20.0 mL of ICG methanol solution (0.5 mg/mL) and were stirred at 25 °C for 24 h in the dark. Following the termination of the reaction, the combined solution was centrifuged, washed with methanol, and the absorbance of the supernatant at 775 nm was measured. The HMTNs@PMO-DOX/ICG was acquired after drying under vacuum at 40 °C for 24 h, and the loading rate of ICG was determined according to formula (1).

0.030 g of HA [36] 0.030 g of NHS and 0.030 g of EDC were dissolved in 30.0 mL of PBS buffer (pH = 7.4), and the reaction was stirred at 25 °C for 12 h in the dark. Subsequently, 0.030 g of HMTNs@PMO-DOX/ICG were dispersed in the above HA-containing solution, and the reaction was carried out at 25 °C for 24 h. To acquire HMTNs@PMO-DOX/ICG@HA, the mixture was centrifuged, washed and dried under vacuum at 30 °C for 24 h.

2.6 Drug release and kinetics study in vitro

5 mg of HMTNs@PMO-DOX/ICG@HA were placed in a dialysis bag (molecular weight cut-off 8000–14000), and 3.0 mL of PBS buffer having different pH (pH = 7.4/5.0) was added to the bag, respectively. The bag was then placed in a beaker with 30 mL of PBS buffer that had the same pH and stirred uniformly at 37 °C. The quantitative drug dissolution was aspirated at a specific time point and then detected its absorbance at 480 nm. The cumulative drug release was determined according to formula (2) [37]

$${\text{ }}C_{c} = C_{t} + \frac{v}{V}\sum\limits_{0}^{{t - 1}} {C_{t} } {\text{ }}$$
(2)

In the formula: CC: actual concentration of DOX released at time t, mg/mL; Ct: apparent concentration of the fluid sample released at time t, mg/mL; v: volume of the sample collected at a predetermined time interval, mL; V: total volume of the sample, mL.

To investigate the drug release behavior of HMTNs@PMO-DOX/ICG@HA in the tumor microenvironment, kinetic fitting of the release behavior in the vitro of drug-composite nanoparticles was carried out using zero-order, first-order, Higuchi, Hixson-Crowell, and Korsmeyer-Peppas models [38,39,40].

Zero order:

$$W = K_{0} t{\text{ }}$$
(3)

First order:

$${\text{ }}\log (100 - W) = \log 100 - {\text{ }}K_{1} t{\text{ }}$$
(4)

Higuchi kinetics:

$$W = K_{H} t^{{\frac{1}{2}}}$$
(5)

Hixson- Crowell kinetics :

$$\left( {100 - W} \right)^{{1/3}} = 100^{{1/3}} - K_{{HC}} t{\text{ }}$$
(6)

Korsmeyer-Peppas equation :

$${\text{ }}M_{t} /M_{\infty} = Kt^{n}$$
(7)

In the formula, K0, K1, KH, KHC, and K represent the zero-order, first-order, Higuchi, Hixson-Crowell and Korsmeyer-Peppas model drug release rate constants respectively, W is the percent drug release at time t, and Mt/M∞ is the fraction of the drug that is released in the dissolution medium.

2.7 Photothermal effect evaluation

Different concentrations of HMTNs@PMO-DOX/ICG@HA solutions (50–1000 µg/mL) were firstly configured with PBS buffer, followed by continuous irradiation with 808 nm near-infrared laser possessing different laser densities (1 W/cm2, 1.5 W/cm2, and 2.0 W/cm2) for 10 min, and the temperature changes of the samples were measured every 2 min by an alcohol thermometer to investigate the correlation between the temperature change and the sample concentration and laser density. Next, a sample with a concentration of 500 µg/mL was continuously irradiated with 1.5 W/cm2 of near-infrared light for 10 min to increase the temperature, and then spontaneously cooled down to room temperature, and this five heating-cooling cycles were used to explore the photothermal stability of HMTNs@PMO-DOX/ICG@HA. A near-infrared imager was used to detect the temperature image changes of HMTNs@PMO-DOX/ICG@HA (500 µg/mL) before and after the irradiation of NIR (808 nm, 1.5 W/cm2, 10 min), and it was compared with PBS aqueous solution. This was done in order to more intuitively observe the temperature changes produced by HMTNs@PMO-DOX/ICG@HA under the stimulation of 808 nm excitation light.

2.8 Biocompatibility assessment

Hemolysis test was utilized to assess the biocompatibility of HMTNs@PMO and HMTNs@PMO@HA. The steps were as follows: the sample solution (HMTNs@PMO/HMTNs@PMO@HA, 50–800 µg/mL) was co-incubated with PBS solution containing mouse erythrocytes in the dark at room temperature for 3 h. Water and pure PBS were employed as the negative and positive controls in this instance, respectively. Centrifugation (10,000 rpm, 5 min) was performed, the supernatant was collected, and its absorbance at 540 nm was measured to determine the hemolysis rate according to formula (8).

$${\text{Hemolysis}}\,{\text{rate}}\left( \% \right){\text{ }} = {\text{ }}\frac{{{\text{Sample}}\,{\text{ absorption}} - {\text{negative }}\,{\text{control }}\,{\text{absorption}}}}{{{\text{Positive}}\,{\text{ control }}\,{\text{uptake}} - {\text{negative}}\,{\text{ control }}\,{\text{absorption}}}} \times 100\%$$
(8)

2.9 Cytotoxicity study

The CCK-8 test was utilized to evaluate changes in the viability of MCF-7 cells before and after selective NIR light irradiation. The steps were as follows: MCF-7 cells in the logarithmic growth phase were collected, counted, and corrected for concentration before being seeded at a density of 8 × 103 cells per well in 96-well plates. The cells were then cultured overnight in an incubator with 5% CO2 at a constant temperature. Subsequently, HMTNs@PMO@HA, HMTNs@PMO-DOX/ICG and HMTNs@PMO-DOX/ICG@HA were co-incubated with the cells for 24 h. After that, the NIR group received NIR laser treatment (808 nm, 1.5 W/cm2, 10 min), and continued to be incubated for 8 h. The medium was removed, and each well was washed with PBS three times. The cells were added to the medium containing 10% CCK-8 according to 150 µL/well and incubated for 2 h in a constant temperature incubator at 37 °C with 5% CO2. Transfer the supernatant 120 µL/well to a new 96-well plate, where an enzyme labeler recorded the absorbance at 450 nm. The relative cell viability value was calculated using formula (9).

$${\text{Relative viability percent}}{\mkern 1mu} = {\mkern 1mu} \frac{{{\text{OD value of experimental group }} - {\text{ OD value of background}}}}{{{\text{Mean OD value of control group }} - {\text{ OD value of background}}}} \times {\text{100}}\%$$
(9)

In the formula, the OD value of background is the absorbance of the multi-well plate whose wells without medium and cells added.

2.10 Cellular uptake

Confocal laser scanning microscopy (CLSM) was utilized to examine the distribution of the samples within MCF-7 cells to look into the tumor-targeting capabilities of the drug-composite nanoparticles. The steps were as follows: MCF-7 cells were co-incubated with HMTNs@PMO-DOX/ICG and HMTNs@PMO-DOX/ICG@HA for 12 h respectively after incubating at 5% CO2, 37 °C for 12 h. The cells were washed with PBS before being stained with DAPI for 30 min, and the distribution of the drug DOX in the MCF-7 cells was observed by CLSM.

2.11 Apoptosis

Flow cytometry was utilized to detect apoptotic changes in MCF-7 cells before and after the introduction of NIR light irradiation after co-incubation with sample HMTNs@PMO-DOX/ICG@HA. MCF-7 cells in the growth phase (2 × 105/well, 6 wells) were taken and incubated in the incubator overnight at 5% CO2, 37 °C. After adding 50 µg/mL of HMTNs@PMO-DOX/ICG@HA solution, the cells were incubated for 48 h. The HMTNs@PMO-DOX/ICG@HA+NIR group was received NIR photostimulation (808 nm, 1.5 W/cm2, 10 min) and the apoptotic changes were detected after another 8 h.

3 Results and discussion

3.1 Synthesis and characterization of nanoparticles

The nanoparticles used silica as the core, and they were coated with a layer of titanium dioxide on the surface through the sol-gel method and finally coated silica on its base to protect calcination and alkali etching to obtain HMTNs. In order to improve the biocompatibility, PMO was coated on the surface of HMTNs to obtain HMTNs@PMO. The morphology of HMTNs and HMTNs@PMO was observed by scanning electron microscopy (SEM) and transmission electron microscopy (TEM). Figure 1a and b show that the prepared HMTNs are spherical nanoparticles with clearly visible hollow structures, and the particle size is mainly distributed about 191.99 nm as depicted in Fig. 1e. Figure 1c and d show that the HMTNs@PMO coated with PMO shell also maintains a relatively uniform spherical structure, with the average particle size distributed about 201.19 nm. Through the comparison of the particle sizes, the particle size after coating is larger than that before coating, indicating the successful wrapping of mesoporous organosilicon layer on the surface of HMTNs.

The microscopic composition of HMTNs@PMO was further analyzed by elemental mapping images. As shown in Fig. 1g, the colors of green, orange, blue, and yellow stand for the distributions of Ti, O, C, and Si, respectively, and the distributions of these four elements all exhibit a regular spherical structure. The color distribution of Ti elements in the outer layer is more concentrated than that in the inner layer, which confirms the successful construction of the hollow structure. The yellow distribution range of Si elements is slightly larger than the green range of Ti elements due to the uniform encapsulation of PMO on the outer surface of HMTNs. The results of elemental mapping images indicated the successful preparation of HMTNs@PMO composite nanoparticles.

The Zeta potential was used to further analyze and confirm whether the prepared nanocomposites were successful. Figure 1h demonstrates the potential changes before and after nanomaterial modification, and the potential of HMTNs shows a negative potential (− 34.49 mV ) [41]. The zeta potential of HMTNs@PMO also shows a negative potential (− 32.56 mV), which is because of the bare silanol hydroxyl groups on the surface of PMO [42]. After APTES modification to introduce positively charged –NH2, the potential of HMTNs@PMO-NH2 increases to a positive potential (+ 3.78 mV) due to the deprotonation of –NH2. After modification by hyaluronic acid, the potential of HMTNs@PMO@HA finally shows a negative potential (− 35.89 mV) due to the strong negative charge effect of the carboxylate ionization. The potential changes throughout the synthesis further confirmed the successful preparation of the nanocomposites.

Fig. 1
figure 1

Scanning electron microscope images (a and c, the magnification in the SEM is ×30,000 for a and ×70,000 for c, respectively.), transmission electron microscope images (b and d) and particle size distribution images (e and f) of HMTNs and HMTNs@PMO, and element mapping of HMTNs@PMO (g), zeta potential analysis of HMTNs, HMTNs@PMO, HMTNs@PMO-NH2 and HMTNs@PMO@HA(h)

Full size image

The crystal structure characteristics of HMTNs and HMTNs@PMO were analyzed by X-ray diffraction (XRD). As depicted in Fig. 2a, there is a broader bun peak appearing around 2θ = 25°. Overall, there is no crystal plane diffraction peak of the standard crystal form of titanium dioxide [43], indicating that the prepared HMTNs and HMTNs@PMO are amorphous structures. From the nitrogen adsorption–desorption curves [44] in Fig. 2b, It is apparent that both HMTNs and HMTNs@PMO exhibit obvious type IV isotherms, which is a typical adsorption isotherm type of mesoporous materials. They both show typical H4-type hysteresis loops when the relative pressure values are 0.5–1.0, which indicates the existence of the mesoporous structures of HMTNs and HMTNs@PMO. Figure 2c and d are the pore size distribution diagrams of HMTNs and HMTNs@PMO, with pore sizes of 3.40 and 3.72 nm, respectively. Table 1 shows the relevant parameters obtained from the BET tests of HMTNs and HMTNs@PMO. It can be found that the specific surface area and pore volume of HMTNs@PMO are smaller than those of HMTNs, which is due to the “blockage” of some of the pore structures on the surface of HMTNs caused by the outer layer PMO, which prevents the nitrogen molecules from entering into the HMTNs completely, resulting in the reduction of nitrogen adsorption-desorption and the decline in specific surface area and pore volume of the composites.

Fig. 2
figure 2

XRD analysis (a) and nitrogen adsorption-desorption curve (b) of HMTNs and HMTNs@PMO, pore size distribution of HMTNs (c) and pore size distribution of HMTNs@PMO (d)

Full size image
Table 1 Specific surface area, pore volume, and pore size of HMTNs and HMTNs@PMO
Full size table

Figure 3 shows the biocompatibility of the nanocarriers explored by hemolysis experiments, and the hemolysis rates of both HMTNs@PMO and HMTNs@PMO@HA are low in the range of 50–800 µg/mL. At a sample concentration of 800 µg/mL, the hemolysis rate of HMTNs@PMO is 2.24%, and the hemolysis rate of HMTNs@PMO@HA is 1.78%. The hemolysis rate of HMTNs@PMO before modification is higher than that of HMTNs@PMO@HA, which is attributed to the existence of bare silanol hydroxyl groups in the PMO shell. There is a direct influence on the relationship between silanol hydroxyl groups and hemolysis rate. Firstly, the exposed silanol hydroxyl groups of PMO can induce the generation of reactive oxygen species and damage the cell structure [45]; secondly, the strong electrostatic force between the trimethylammonium head group (positively charged) in the phospholipid layer of the lipid cell membrane and the silanol hydroxyl groups (negatively charged) causes the damage of the cell structure, which increases the hemolysis rate [46, 47]. After hyaluronic acid modification, the HMTNs@PMO@HA hemolysis rate is reduced due to the modifying effect of hyaluronic acid “shielded” the effect of silanol hydroxyl groups in the PMO shell layer [36]. Hemolysis assays demonstrated the prepared nanocarriers’ excellent biocompatibility.

Fig. 3
figure 3

Hemolysis assay analysis of HMTNs@PMO, HMTNs@PMO@HA

Full size image

3.2 Assessment of photothermal effects

The photothermal conversion experiments were performed to evaluate the photothermal conversion ability of HMTNs@PMO-DOX/ICG@HA which used ICG as a photothermal agent. The assays were carried out to explore the relationship between temperature and sample concentration, NIR optical power density, and evaluate the photothermal stability of the particles. Figure 4 a–e show the temperature changes of HMTNs@PMO-DOX/ICG@HA (50–1000 µg/mL) under different power densities of 808 nm NIR light. These figures depict that the temperature changes of any concentration of HMTNs@PMO-DOX/ICG@HA increase with the 808 nm NIR light power density increases, showing a power density-dependent change. Figure 4f shows the temperature changes of different concentrations of HMTNs@PMO-DOX/ICG@HA under 808 nm NIR light irradiation at the same power density (1.5 W/cm2 for 10 min). It can be seen from the temperature change curves in the figure that the temperature increases with the increase of the sample concentration under irradiation at the same excitation light intensity, presenting a sample concentration-dependent change. When the continuous irradiation time is 4 min, the temperature of HMTNs@PMO-DOX/ICG@HA basically elevates to the highest temperature, which indicates that the drug delivery system using ICG as a photothermal agent could rapidly elevate the temperature to the highest temperature in a short period, enabling the tumor cells to achieve thermal ablation. Figure 4g uses 808 nm NIR light to stimulate heating and then natural cooling to explore the photothermal stability of HMTNs@PMO-DOX/ICG@HA (500 µg/mL, 1.5 W/cm2) in five cycles. As can be seen from the figure, it has the highest photothermal conversion activity in the first temperature cycle, which causes a sharp increase in temperature in a short period. In each subsequent heating process, the maximum temperature of HMTNs@PMO-DOX/ICG@HA that can be achieved gradually decreases, which is because more heat generates in the solution in the previous cycle, and in this hot environment, the structure of the photothermal agent ICG is destroyed [48], which reduces the photothermal conversion ability. As a result, the temperature in the next cycle cannot return to the highest temperature reached in the previous cycle. Therefore, due to ICG’s own “photobleaching” and poor thermal stability, HMTNs@PMO-DOX/ICG@HA was more suitable for treatment modalities that required rapid heating in a short period. In order to observe the photothermal conversion effect of HMTNs@PMO-DOX/ICG@HA more intuitively, Fig. 4h utilized NIR photoimager to detect the temperature changes of PBS and HMTNs@PMO-DOX/ICG@HA before and after NIR light irradiation. From the color change process in the figure, it can be observed intuitively that the temperature of HMTNs@PMO-DOX/ICG@HA is dramatically increased after NIR light irradiation, while PBS doesn’t produce obvious changes during this process. The above results further indicated that HMTNs@PMO-DOX/ICG@HA can realize the combined treatment of chemotherapy and photothermal therapy by co-delivery of chemotherapeutic drugs and photothermal agents.

Fig. 4
figure 4

Temperature changes of HMTNs@PMO-DOX/ICG@HA under 808 nm NIR light irradiation with different power densities (a–e), temperature changes of HMTNs@PMO-DOX/ICG@HA with different concentrations under 808 nm NIR light irradiation (1.5 W/cm2 for 10 min) (f), photo-thermal stability investigation (g) and photothermal imaging analysis (h)

Full size image

3.3 Analysis of drug loading and in vitro drug release behavior

Drug loading experiments demonstrated that the loading rates of the chemotherapy drug DOX and the photothermal agent ICG in HMTNs@PMO-DOX/ICG@HA were 40.21% and 25.78% respectively, indicating that the drug carrier had a high drug loading rate. Figure 5a simulates the release behavior of drug carriers in different pH environments, and it is discovered that there are considerable discrepancies, Fig. 5b further validates the significant difference in drug release from drug carriers in different pH environments. In the condition of pH = 7.4, the release rate of DOX from this drug carrier is slow, and its cumulative drug release amount is 4.63% for up to 240 h. This is due to the fact that in a neutral environment, it is difficult to cause damage to the hyaluronic acid on the surface of the HMTNs@PMO-DOX/ICG@HA, making it impossible to open the pores in the drug delivery system, which in turn exhibits a slow release rate and low cumulative drug release amount from DOX [36]. In the condition of pH = 5.0, the release rate of DOX and the cumulative drug release amount in the drug carrier increase significantly, which is attributed to the acidic condition at low pH destroying the amide bond formed between hyaluronic acid and PMO in the system. It allows the release of DOX through the pore structure in the drug carrier, which shows an accelerated release rate and an increase in the cumulative drug release amount. On the whole, the cumulative release of the drug tends to level off, so the system is more suitable for the long-term slow release of the drug.

The pH range of normal human plasma is between 7.35 and 7.45 [49]. To indirectly verify the stability of HMTNs@PMO-DOX/ICG@HA in plasma, the PBS buffer solution with pH = 7.4 was used to simulate the physiological environment of the human body to study the stability of HMTNs@PMO-DOX/ICG@HA. The method was as follows: 5 mg of HMTNs@PMO-DOX/ICG@HA were placed in a beaker containing 20 mL of PBS buffer with pH = 7.4, and stirred uniformly at 37 °C, and the drug composite nanoparticles were obtained by centrifugation after 240 h. The morphology and size changes of HMTNs@PMO-DOX/ICG@HA before and after stirring in PBS buffer for 240 h were examined by scanning electron microscopy to explore its stability. The results are shown in Fig. 5c–f. The morphology of the drug composite nanoparticles has basically not changed before and after stirring in PBS buffer with pH = 7.4 for 240 h, and their size has also basically not changed, proving that HMTNs@PMO-DOX/ICG@HA is stable in PBS buffer solution with pH = 7.4, which indirectly proves that it is relatively stable in plasma.

The kinetic fitting analysis of DOX release behavior in HMTNs@PMO-DOX/ICG@HA was performed by utilizing zero-order, first-order, Higuchi, Hixson-Crowell and Korsmeyer-Peppas models. Figure 5g–k and Table 2 show the results of the fitted data. It can be observed that the system has good linear relationships in the five kinetic analysis models, and the R2 values are all greater than 0.9. Among them, the release behavior of HMTNs@PMO-DOX/ICG@HA has the highest degree of fit to the Higuchi model with the R2 as high as 0.997. The model of Higuchi is mainly utilized to describe drug delivery systems that use dissolution and diffusion as the release mechanism. At pH = 7.4, the gating effect of hyaluronic acid results in slow infiltration of PBS buffer and increases resistance to drug release. At pH = 5.0, the amide bond breaks, and the gating opens, allowing the solvent to penetrate through the pores, creating a drug concentration gradient and increasing drug release [38,39,40]

Fig. 5
figure 5

HMTNs@PMO-DOX/ICG@HA drug release behavior in vitro (a and b, ***P < 0.001). Scanning electron microscope images (c and d, the magnification in the SEM is × 70,000 for c and × 100,000 for d, respectively.) and particle size distribution images (e and f) of HMTNs@PMO-DOX/ICG@HA before and after stirring in PBS buffer solution at pH = 7.4 for 240 h. And models of HMTNs@PMO-DOX/ICG@HA drug release in zero-order (g), first-order (h), Higuchi (i), Hixson-Crowell (j), and Korsmeyer-Peppas (k)

Full size image
Table 2 Fitting parameters of HMTNs@PMO-DOX/ICG@HA
Full size table

3.4 Analysis of cytotoxicity

The CCK-8 test was utilized to assess the toxic effects of HMTNs@PMO-DOX/ICG@HA on MCF-7 cells to further explore the tumor cell-inhibitory properties of the drug-composite nanoparticles. According to Fig. 6a, HMTNs@PMO@HA has high cell viability values in the concentration scale of 0.1–100 µg/mL, indicating that it does not significantly cause cytotoxicity and has the potential to be a biologically sound drug carrier. In Fig. 6b, it can be seen that pure NIR light and HMTNs@PMO@HA (100 µg/mL) possess high cell viability values, whereas as the antitumor drug DOX and the photosensitizer ICG have cell viability values of 5% and 72.77% at a concentration of 100 µg/mL respectively. The significant discrepancy in cell viability between HMTNs@PMO-DOX/ICG and HMTNs@PMO-DOX/ICG@HA in the graph (83.31% and 48.85%), and also between the two after the introduction of NIR light (81.26% and 27.68%) is attributed to the fact that functionalized modification of the surface of HMTNs@PMO with HA enhances the targeting of the system to MCF-7 cells, which specifically binds to the CD44 receptor overindulgently expressed on the surface of MCF-7 cells and in turn improves the internalization of the cells to the drug-composite nanoparticles, releasing DOX and inducing apoptosis [28]. Compared with HMTNs@PMO-DOX/ICG@HA, HMTNs@PMO-DOX/ICG@HA+NIR exhibits a more significant cytotoxic effect, which is attributed to the introduction of NIR light. When the drug carrier targets and enters into the MCF-7 cells, ICG successfully converts the absorbed light energy into heat energy under the stimulation of NIR light, inducing cell apoptosis in a thermal environment. At the same time, the released DOX induces tumor cells to die abnormally under the mediation of chemotherapeutic drugs, exerting the role of photothermal synergistic chemotherapy to inhibit tumor cells. The results revealed the importance of HA targeting and photothermal conversion in the prepared drug composite nanoparticles

Fig. 6
figure 6

a Analysis of cell viability of HMTNs@PMO@HA. b Cell viability analysis of NIR, DOX, ICG, HMTNs@PMO@HA, HMTNs@PMO-DOX/ICG, HMTNs@PMO-DOX/ICG@HA, HMTNs@PMO-DOX/ICG+NIR and HMTNs@PMO-DOX/ICG@HA+NIR (100 µg/mL) co-incubated for 12 h with MCF-7 cells, ****p < 0.0001

Full size image

3.5 Analysis of the cellular uptake assay

In order to further more intuitively explore the tumor-targeting properties of hyaluronic acid on the drug-composite nanoparticles, MCF-7 cells were chosen as the tumor cell model and confocal laser scanning microscopy (CLSM) was utilized to observe the uptake of HMTNs@PMO-DOX/ICG and HMTNs@PMO-DOX/ICG@HA by MCF-7 cells, respectively. It is apparent from Fig. 7 that HMTNs@PMO-DOX/ICG@HA modified by functionalizing hyaluronic acid shows obvious red fluorescence (the red fluorescence of DOX itself), and the fluorescence intensity is significantly higher than that of HMTNs@PMO-DOX/ICG without hyaluronic acid modification, which is attributed to the fact that the drug-composite nanoparticles modified by hyaluronic acid could proactively target to MCF-7 cells and enhance the cellular uptake of HMTNs@PMO-DOX/ICG@HA. The gating is opened and DOX is released from the pore of the drug carrier to exert chemotherapeutic effects under the condition of tumor microacidity after HMTNs@PMO-DOX/ICG@HA are selectively swallowed into the MCF-7 cells. The results further demonstrated that HMTNs@PMO-DOX/ICG@HA modified by hyaluronic acid exhibited a clear tumor-targeting effect, reflecting the successful construction of the tumor-targeting drug delivery system.

Fig. 7
figure 7

Confocal scanning microscopy imaging analysis of HMTNs@PMO-DOX/ICG and HMTNs@PMO-DOX/ICG@HA after 12 h co-incubation with MCF-7 cells

Full size image

3.6 Analysis of apoptosis

Flow cytometry was utilized to explore the apoptotic changes of MCF-7 cells after 48 h of co-culture with HMTNs@PMO-DOX/ICG@HA and HMTNs@PMO-DOX/ICG@HA+NIR. As can be observed from Fig. 8, the apoptosis rate of HMTNs@PMO-DOX/ICG@HA is 33.8%, and HMTNs@PMO-DOX/ICG@HA+NIR is 43.4%. The apoptotic effect possessed by the former originates from the chemotherapeutic effect produced by DOX released from HMTNs@PMO-DOX/ICG@HA in MCF-7 cells, which has an obvious effect of inducing apoptosis in tumor cells. The apoptosis rate of HMTNs@PMO-DOX/ICG@HA+NIR increases to 43.4% under the external stimulation of NIR light, which is 9.6% higher than that of HMTNs@PMO-DOX/ICG@HA without NIR light irradiation. When NIR light irradiates the MCF-7 cells, the ICG exerts the photothermal conversion property of photothermal agents to elevate the local temperature of tumor cells, which makes the tumor cells undergo thermal ablation and promotes apoptosis, resulting in photothermal efficacy. Under the combined effect of chemo-photothermal therapy, HMTNs@PMO-DOX/ICG@HA+NIR had a notable apoptosis-inducing effect on MCF-7 cells, and the efficacy of this combined effect was higher than that of single chemotherapy.

Fig. 8
figure 8

Flow cytometry analysis of apoptosis in HMTNs@PMO-DOX/ICG@HA and HMTNs@PMO-DOX/ICG@HA+NIR treated MCF-7 cells. Q1: necrotic cells, Q2: late apoptotic cells, Q3: early apoptotic cells, Q4: living cells

Full size image

3.7 Future work

There are still some limitations in the current work, and only basic in vitro studies have proven the safety, targeting and effectiveness of the drug delivery system. In future work, relevant in vitro experiments will be used to further verify the feasibility and effectiveness of the drug delivery system constructed in this study in terms of cell toxicity, targeting efficacy under dynamic conditions, penetration and ablation rate in different organs.

4 Conclusion

In this paper, the chemical/photothermal platform, HMTNs@PMO-DOX/ICG@HA, was successfully constructed, which had a high drug loading rate for DOX and ICG (DOX: 40.21%; ICG: 25.78%). Hemolysis and cytotoxicity assays demonstrated that HMTNs@PMO@HA did not have significant cytotoxic effects and had the potential to be a biologically sound drug carrier. In vitro drug release probes showed that the platform exhibited pH-stimulated responsive drug release and was more suitable for long-term sustained release of chemotherapeutic drugs. The photothermal effect evaluation results showed that HMTNs@PMO-DOX/ICG@HA possessed excellent photothermal conversion ability. The cellular experiments confirmed the successful construction of the HMTNs@PMO-DOX/ICG@HA chemo-photothermal combined treatment system. After introducing the NIR light, the system had an effect on MCF-7 cells which reflected by a significant decrease in cell viability value from 48.85 to 27.68%, and a significant increase in apoptosis rate from 33.8 to 43.4%. The above results demonstrated the successful construction of the chemo/photothermal treatment platform, which released the photothermal agent ICG and chemotherapeutic agent DOX after HA targeting into the tumor cells to locally elevate the tumor temperature and cause thermal damage through near-infrared light, coupled with the direct killing effect of the chemotherapeutic agent DOX, to achieve the combined chemo/photothermal treatment and thereby enhanced anti-tumor effect.