DEVELOPMENT OF COPPER SULFIDE NANOPARTICLES FOR PHOTOTHERMAL AND CHEMO THERAPY OF CANCER CELLS

Cancer has become number one cause of death. Conventional treatment includes surgery, chemotherapy, radiation therapy, or combination. The combinatorial therapy in one system is highly efficient and economical. Herein, two drug delivery systems with chemo and photothermal therapy are developed in order to enhance the therapeutic efficacy in A549 human lung cancer cells. The first one is hollow copper sulfide nanoparticle carrying doxorubicin (PEG-HCuSNPs-DOX). The second one is mesoporous silica coated CuS nanoparticles (CuS NPs) loaded with doxorubicin (DOX) (PEG-CuS@MSNs-DOX). Both nano-drug delivery systems are pH sensitive, laser responsive, and photothermal convertible. CuS NPs are the photothermal sensitizers in both drug delivery systems. However the drug loading efficiency is much higher in the PEG-CuS@MSNs-DOX, whereas its drug release rate is much slower. In addition, the anti-cancer efficacy of PEG-HCuSNPs-DOX is higher than that of PEG-CuS@MSNs-DOX. Five chapters are prepared in this thesis. Each chapter includes an independent manuscript and separate abstract. Chapter 2 serves as preparation experiment for chapter 3. Chapter 4 is a review that expands the application of CuS NPs to transdermal delivery. Chapter 5 is a mini review on the in vivo application of CuS NPs 1. Cancer Photothermal Chemo Therapy Using Hollow Copper Sulfide Doxorubicin	  Nanoparticles The design and synthesis of the multifunctional nanoparticles responsive to external stimuli provides potential applications in biomedical fields such as controlled drug delivery. Here, near infrared (NIR) laser-controlled fast and effective tumor cell killing is achieved based on the pH sensitive and NIR light responsive hollow copper sulfide nanoparticles chelated with doxorubicin molecules (HCuSNPs-DOX). Laser exposure at 900 nm and acidic environment facilitate the release of DOX from HCuSNPs-DOX. Spontaneously, the released DOX forms DOX/Cu complex and generates cell-killing reactive oxygen species. Laser exposure to HCuSNPs-DOX also disrupts the integrity of the cell membrane instantly. The IC50 of HCuSNPs-DOX with and without laser treatment was 4.0 and 7.6 μg/mL CuS, respectively. The approach developed here offers compelling chances for quick-responsive anticancer therapy. 2. Facile Direct Dry Grinding Synthesis of Monodisperse Lipophilic CuS Nanoparticles Copper sulfide with near-infrared light absorption property is recently attracting broad interest as a photothermal carrier for smart cancer therapy. Lipophilic copper sulfide nanoparticle is preferred for high performance biomedical applications due to the high affinity with tissues. But it requires complex multi-step synthetic process under severe condition. Here, synthesis of hydrophobic copper sulfide possessing surface plasmon resonance was retained by direct dry grinding of copper(II) acetylacetonate with sulfur under ambient environment. The formed CuS nanoparticles were in uniform size of ~10 nm, and they were monodispersed in pure chloroform. Each covellite CuS nanocrystal surface was modified with oleylamine through hydrogen bonding between sulfur atoms and amine groups of oleylamine. While those oleylamine capped CuS nanoparticles showed uniform morphological features, they demonstrated near-infrared light absorption for photothermal applications. The facile and mild synthetic methodology described here opened a powerful pathway for the design and preparation of photothermal lipophilic copper sulfide nanomaterials for smart cancer therapy. 3. Multifunctional Mesoporous Silica-Coated CuS Nanoparticles for Cancer Therapy: Synthesis, Characterization and in vitro Evaluation Chemo therapeutic drug-caused side effects are commonly seen in clinical practice due to nonspecific toxicity and low therapeutic efficiency. Herein, we reported a cancer chemo-photothermal multifunctional drug delivery system. Polyethylene glycol decorated mesoporous silica nanoparticles entrapping CuS nanoparticles (PEG-CuS@MSNs) were successfully synthesized and characterized for the drug delivery application. Doxorubicin (DOX)-loaded PEG-CuS@MSNs showed laser stimulated and pH-responsive properties. In vitro cell experiments demonstrated that DOX-loaded PEG-CuS@MSNs combining laser exposure achieved the highest rate of death of A549 cells, in comparison to that of PEG-CuS@MSNs-DOX chemotherapy alone. These findings provided a promising drug delivery system for cancer combinatorial therapy, which could significantly reduce drug dose and improve patient compliance. 4. Laser ablation-enhanced transdermal drug delivery Transdermal delivery offers an excellent route for drug and vaccine administration. Nonetheless, it presents a critical challenge due to the skin’s lipid-rich outer stratum corneum layer. Laser ablation perforates epidermis through selective photothermolysis, making skin more permeable to hydrophilic and macromolecular drugs such as peptides, proteins, and genes. This review summarizes recent applications to laser ablation-enhanced transdermal delivery. Needleand pain-free transcutaneous drug delivery via laser ablation provides an alternative approach to achieve local or systemic therapeutics. 5. Cancer Photothermal Therapy and CuS Nanoparticles This manuscript is being prepared according to the format of Lasers in Medical Science as a review article.


INTRODUCTION
Cancer has nowadays become one of the most deadly diseases in the world. As reported in 2013, the 5-year global cancer prevalence is estimated to be 28. 8  Copper sulfide nanoparticles are a new class of photothermal sensitizer providing an affordable counterpart for gold nanoparticles. The light absorption of the former is affected by the surrounding environment. 6,7 Originating from the d-d* transition of Cu 2+ ions in copper sulfide, such nanoparticle exhibits stable light absorption towards near-infrared (NIR) light irradiation (650-900 nm), 8 which can penetrate through normal tissues with minimal thermal injury. 9 Instantaneously upon NIR light absorption, copper sulfide nanoparticles generated heat and photothermally ablated tumor in vivo after intratumor 10,11 or intravenous injection. 12 In our previous work, hollow copper sulfide nanoparticles (HCuSNPs) were applied for photothermal ablation-enhanced transdermal drug delivery. 13

RESULTS AND DISCUSSION
The preparation of PEG-HCuSNPs-DOX was carried out as illustrated in Scheme 1.

Scheme 1.
Schematic illustration of the preparation procedure of the PEG-HCuSNPs-DOX and NIR laser controlled drug release process.
The transmission electron microscopy (TEM) image of the as-prepared HCuSNPs demonstrated hollow structures with the average diameter of 75 ± 11 nm ( Figure 1a).
The shells were ~20-nm thick, and consisted of 8-12-nm large nanoparticles. After surface modification with thiolated PEG, a thin layer (thickness ~4 nm) was clearly observed on the particle surfaces, while the initial structures of HCuSNPs did not change ( Figure 1b). Loading of DOX did not significantly change the size and morphology of the nanoparticles (Figure 1c). Dynamic light scattering (DLS) analysis revealed that the hydrodynamic particle size of the PEG-HCuSNPs-DOX was 80 ± 10 nm, which agreed well with the TEM observations. As shown in the UV-Vis spectrum ( Figure 1d), PEG-HCuSNPs exhibited strong absorbance peak centered at ~1050 nm, which was ascribed to the d-d* transition of copper sulfide. 20 Similar absorbance was found in the UV-Vis spectrum of PEG-HCuSNPs-DOX, but with a minor red shift of 5 nm due to the DOX chelation on HCuSNPs. 15 When compared with the green aqueous dispersion of PEG-HCuSNPs, the aqueous dispersion of PEG-HCuSNPs-DOX was more brownish (Figure 1e). The drug loading increased with the increase of time (Figure 1f). The PEG-HCuSNPs and PEG-HCuSNPs-DOX remained stable in DI water at room temperature for at least 3 months. The DOX loading efficiency of the PEG-HCuSNPs-DOX was optimized ~6.0 wt.%.  and fluorescence data clearly indicating the existence of copper and doxorubicin. At pH 5, the band intensity of DOX near 478 and 498 nm decrease upon the partial formation of DOX/Cu 2+ 2:1 complex. In addition, the absorbance peak near 535 nm increased ( Figure 3a). The released samples with or without laser treatment all showed similar absorbance curves to the complex rather than free DOX, indicating formation of the DOX/Cu 2+ complex. When pH increased to 6, both 2:1 and 1:1 complex exist.
As shown in Figure 3a, the absorbance of the DOX/Cu 2+ complex at 480 nm continued to decrease and the broad band was centered at 506 nm. The band at 550 nm increased significantly. The released samples showed similar changes.
In comparison with free DOX, the CD spectrum of DOX/Cu 2+ at pH 5 and 6 showed negative and positive bands at 490 and 550 nm respectively, which were typical of DOX/Cu 2:1 complexes (Figure 3b)   DOX induced the generation of reactive oxygen species (ROS), thus causing the oxidation of lipid, protein, and DNA in cancer cells [20][21][22][23][24][25] . More interestingly, transition metal ion has been indicated to be a critical cofactor facilitating this process. 15 To prove that DOX/Cu 2+ complex released from PEG-HCuSNPs-DOX could assist ROS generation and therefore induce therapeutic effects on cancer cells, ROS level was tested via fluorescence imaging and quantification. As shown in the fluorescent microscope images (Figure 6a), only weak signal of ROS species was detected from free DOX, PEG-HCuSNPs or PEG-HCuSNPs with NIR laser exposure. However, existence of ROS was clearly shown by the green fluorescence in the cases of PEG-HCuSNPs-DOX and DOX/Cu 2+ complex with NIR laser irradiation ( Figure 6a).
The intensity of reactive oxygen species signal increased after NIR laser irradiation ( Figure 6b), because the HCuSNPs released Cu 2+ ions 13 and facilitated DOX molecules to chelate with Cu 2+ . As comparisons, the generation of reactive oxygen species by hollow gold nanoparticles (HAuNPs) and DOX modified gold nanoparticles (HAuNPs-DOX) was studied under the identical condition. However, no such fluorescence was observed from either sample despite of the NIR laser irradiation ( Figure 6b). Therefore, the cancer killing effect of PEG-HCuSNPs-DOX could be attributed to a combination of DOX release, photothermal effect, and copper induced ROS generation following NIR laser irradiation. The One-way analysis of variance (One-way ANOVA) results were shown in Table 1. and Table 2.  The total effect of PEG-HCuSNPs-DOX on cell viability was evaluated using 3-(4,5-dimethylthiazol-2-yl)-2,5-diphenyltetrazolium bromide (MTT) assay. As shown in Figure 7a, NIR laser exposed PEG-HCuSNPs-DOX exhibited the highest

SUMMARY AND CONCLUSIONS
In summary, NIR laser controlled anticancer nanocomposite was achieved by modification of PEG-HCuSNPs with DOX through chelating interaction. The nanocomposite was capable of pH-sensitive drug release, photothermal conversion, and generating ROS in response to NIR laser irradiation. With the NIR laser responsive multi-functions, highly controlled and effective cancer killing performance was achieved. The current work sheds a considerable light on the smart antitumor materials by NIR laser control for transdermal cancer therapy.

Chemicals. Doxorubicin hydrochloride and Dulbecco's Modified Eagle Medium
(DMEM) were purchased from AK scientific and Lonza, respectively.
where E DOX represents the DOX loading efficiency, W DOX stands for the weight of DOX extracted from the nanoparticles, and the W PEG-HCuSNPs-DOX shows the weight of PEG-HCuSNPs added.
In Vitro Release. In vitro release of DOX from the as-prepared nanoparticles by adding 10 mg PEG-HCuSNPs-DOX in a dialysis tube (1000 Da cut off; Sigma-Aldrich), placing the tube in 100 mL of phosphate-buffered saline (PBS, pH 7.4, 10 mM) containing 10% BSA, and shaking constantly at 150 rpm at 37°C. The NIR laser was administered as needed after taking out the sample from the dialysis tube at 2.0 W/cm 2 for 15 s at a time interval of 1 h. The in vitro DOX release was continued for 4 h. The release medium was replaced with a fresh one at a determined interval to maintain sink conditions. The amount of released DOX was monitored by fluorescent measurement. The fluorescent emission peak at λ = 590 nm under excitation light (λ = 485 nm) of each solution was recorded to determine the released DOX amount. As a comparison, in vitro release was proceeded identically without NIR irradiation.All measurements were performed in triplicate.

MTT
Assay. The final concentration of DOX was 0.5 µM, and the amount of PEG-HCuSNPs was kept the same in the wells. After incubating for 4 h, the drug-containing medium was replaced with new one and irradiated with NIR laser (900 nm, 2 W/cm 2 , 15 s) . Then, the medium was substituted with 1 mM DFCH-DA medium and incubated for 30 min.
After washing with PBS for 3 times, the wells were observe by using the fluorescent microscope.
To determine the ROS level in A549 cells, A549 cell line was cultured in the 96-well plate at a concentration of 10,000 per well. Then, the cells were treated identically by replacing the cell culture medium with DMEM, incubated for 3 h, changing the culture medium with pre-warmed freshmedium, adding samples, incubated for 4h, replacing medium with DMEM, exposed with corresponding light irradiations, substituting the medium with 1 mM DFCH-DA medium, incubated for 30 min, washing with PBS for 3 times, adding 100 µL cell lycis buffer. Finally, the treated specimens were subjected to fluorescent test (excitation λ = 480 nm, emission λ = 530 nm) to measure the protein content. As a comparison, DMEM containing 5 mM NAC was applied as the cell culture medium and tested through the same procedure. All data were expressed as mean SD and IC 50 values were calculated by using nonlinear regression analysis. The statistical significance was determined using a t test.
A p value less than 0.05 (i.e., p < 0.05) was considered to indicate statistical significance for all comparisons.   The photothermal conversion effect of the CuS nanoparticles is independent of the surrounding environment. [6,7] These features are especially useful for controlled drug delivery and photothermal cancer therapy. [8][9][10] For the synthesis of CuS nanoparticles with desired nanostructures, a series of approaches have been developed, such as hydrothermal [11], solvothermal method [12], solid-state reaction [13], microemulsion [14], and reflux condensation [15] have been developed. In order to endow the CuS nanoparticles with NIR absorption, the as-prepared nanoparticles are usually further oxidized to produce vacancies in the crystalline structures. [ Recently, lipophilic nanomaterials have been developed for their drug delivery into hydrophobic tissues such as brain and vascular tissues. To retain CuS nanoparticles dispersible in organic phase, hot injection [17], cation exchange, [18] and solventless approach [19] have been reported. Among them, hot injection method based on high temperature reaction of copper (II) acetylacetonate and elemental sulfur or sulfur provider (e.g., dodecanethiol) has been widely used. However, the lipophilic CuS nanoparticles synthesized by these methods are not able to absorb NIR light.
Thus, they require additional complex oxidization treatment to show photothermal performance.
In this paper, lipophilic CuS nanoparticles were synthesized by directly grinding copper (II) acetylacetonate with sulfur in oleylamine at room temperature. Within a few minutes of grinding in the ambient environment, the CuS nanoparticles were attained in high yield. The resulting CuS nanoparticles were in uniform particle size of

RESULTS AND DISCUSSION
The TEM image of the CuS nanoparticles synthesized by the dry grinding process is shown in Fig. 1a. Many nanoparticles were clearly observed. These metallic nanoparticles were mainly in cubic geometry, and they were monodispersed. Some minor aggregation was caused by the evaporation of chloroform component during the TEM sample preparation process, which was a common situation. Calculated based on at least 300 particles, the average size for these CuS nanoparticles was ~10 nm. This result matched well with the hydrodynamic particle diameter of the DLS analysis (Fig.   2), indicating excess surfactant was cleared and monodisperse fine nanoparticles remained. As a comparison, CuS nanoparticles were prepared through the previously reported solution based technique. As shown in TEM (Fig. 1b), the formed CuS nanoparticle size was ~9 nm on average, and they were similar as the above nanoparticles obtained by the dry grinding process. The solution based approach derived nanoparticles were more spherical, because the liquid environment inhibited directional crystal growth of the nanocrystals. Therefore, the dry grinding synthesis approach achieved the fine nanocrystals, which was almost identical to the traditional solution based method.
The XRD pattern of the powder sample prepared by the dry grinding method nm. This size was relatively smaller than the particle size measured in TEM images (10 nm), because a minor amorphous oleylamine layer was modified on the nanocrystal surface. These characteristic peaks were identical to those prepared through the solution based method. Therefore, the current CuS nanoparticles prepared through dry grinding process formed high quality fine covellite CuS nanoparticles.
XPS spectra of the as-prepared CuS nanoparticles were summarized in Fig. 4.
The Cu 2p XPS spectrum exhibited 2p 3/2 peak at 932.0 eV and 2p 1/2 peak at 952.2 eV, which were typical peaks for Cu(II) in copper sulfide. [21] The C 1s peak was resolved as two peaks located at 284.6 eV and 285.7 eV, which respectively corresponded to  (Fig. 5), supporting that the current CuS nanoparticles were capped with oleylamine. Hydrophilic Sulfur atoms in CuS was qualified electron acceptors. [16] Although it hardly interacted with the hydrophobic alkyl terminals of the oleylamine, it readily accepted electron from the amine group in oleylamine, forming S-H bonds.
In the FTIR spectrum of the resultant CuS nanoparticles (Fig 6), the broad band at ~3450 cm -1 was assigned to N-H stretching vibration of the amine group in oleylamine, [24] the two bands at 2922 cm -1 and 2852 cm -1 were assigned to the asymmetric (ν as ) and symmetric (ν s ) stretching vibrations of methylene (CH 2 =CH) in the alkyl chain of oleylamine, the bands centered at 1634 cm -1 was attributed to N-H bending vibration. [25][26][27] All of these characteristic bands were in fair agreement with the FTIR spectrum of pure oleylamine. Hence, oleylamine was capping on the CuS nanoparticles.
Interestingly, the dry grinding synthesized CuS nanoparticles demonstrated broad NIR absorption peaks centered at ~ 1100 nm (Fig. 7), which was very close to the

CONCLUSIONS
Synthesis of monodisperse CuS nanocrystals was achieved by a facile one step dry grinding process. The nanoparticles were composed of covellite phase CuS, and the particle size was finely controlled as ~10 nm. The CuS nanoparticle surface was capped with oleylamine by hydrogen bonding between sulfur atoms with amine group of oleylamine. While the resultant CuS nanoparticles were highly comparable with those prepared through traditional solvothermal method, the current approach was carried out at ambient condition and decreased use of toxic solvents. This environmental benign opened a facile pathway for the large-scale production of photothermal nanocrystals for applications smart drug delivery.
ABSTRACT: Chemo therapeutic drug-caused side effects are commonly seen in clinical practice due to nonspecific toxicity and low therapeutic efficiency. Herein, we reported a cancer chemo-photothermal multifunctional drug delivery system. States. [1] Various methods have been developed in addition to chemotherapy, such as nano technique, targeted delivery, photothermal therapy, photodynamic therapy, etc. [2] One highlighted trend is the application of nanotechnology, which delivers drug more precisely at cancer cells and brings less damage to normal cells, thus diminishing side effects. Beside, photothermal therapy can ablate cancer cells. Combination of chemo and photothermal therapy with targeting feature into a nano delivery system would be a practical and efficient solution worth trying.
Copper sulfide nanoparticles (CuS NPs) are a new class of photothermal sensitizer.
Their light absorption is not affected by the surrounding environment. [3,4] They exhibit stable light absorption towards near-infrared (NIR) light irradiation (650-900 nm), [5] which will bring minimal thermal injury to normal tissues. [6] Immediately upon NIR light absorption, CuS NPs generate heat and photothermally ablate tumor in vivo after intratumor [7,8] or intravenous injection. [9] Although CuS NPs are promising, when applying to drug delivery, the nanoparticle itself has limitation as a platform. [10] Since the surface is only one layer to attach the chemicals, the loading efficiency is limited. To address this problem, mesoporous silica is chosen considering its large pore size and high surface area.
Moreover, the technique of synthesizing a layer of mesoporous silica on the surface of inorganic template is relatively mature. There have been many reports on the preparation of nanocrystals coated with mesoporous shells such as gold nanorods, [11] iron oxide, [12] manganese oxide nanoparticle [13] , graphene nanosheet, [14] etc. However, studies on CuS NPs coated with mesoporous silica have not been reported.
In this study, a chemo photothermal drug system was formulated to increase drug loading efficiency. Specifically, mesoporous silica spheres containing CuS NPs in the core and DOX loaded in the silica channels (PEG-CuS@MSNs-DOX) were prepared.
Furthermore, these mesoporous silica spheres were applied to photothermal therapy. A procedure for the synthesis of the CuS@MSNs-DOX is shown in Scheme 1.

Chemicals. Doxorubicin hydrochloride and Dulbecco's Modified Eagle
Medium (DMEM) were purchased from AK scientific and Lonza, respectively. All the other chemicals were bought from Sigma-Aldrich without further purification. The water was purified by using a Milli-Q Synthesis system (Millipore) with the resistivity higher than 18.2 MΩ·cm.

CuS nanoparticle preparation.
Copper acetylacetonate (0.131 g) and 0.032 g sulfur All data were expressed as mean SD. The statistical significance was determined using a t test. A p value less than 0.05 (i.e., p < 0.05) was considered to indicate statistical significance for all comparisons.

RESULTS AND DISCUSSION
Oleylamine coated CuS NPs was synthesized with a newly developed facile non-solvent method. The TEM image suggested that CuS NPs were in flake shape ( Figure 1a) with an average size of 12 nm. The lateral view of some particles demonstrates thin and rod-like structure, which agreed well with the CuS NPs prepared by the complicated multi-step techniques. [15] It was considered that several CuS nanoflakes stacked and were wrapped together into one silica nanoparticle during the sol-gel reaction (Figure 1b). The average size of CuS@MSN was ~40 nm, meeting the practical requirement for drug delivery. [16] . Figure 1c showed the PEG-CuS@MSNs nanoparticles (right) turned to purple after loading DOX (left). The loading efficiency was as high as 18.4%. DLS measurement presented a particle diameter of 58 nm. This was close to the TEM result, but it was slightly larger (Figure 2). Because DLS measurement acquired hydrodynamic data, and the swollen state of the nanoparticle inevitably became bigger than the value at dry shrunk state. [17] Zeta potential of CuS@MSN was -36.2 mV. Although the initial nanoparticle surfaces are negatively charged, the coating of PEG on the surface brought the zeta potential to nearly neutral, ~0.5 mV, which was preferable for drug delivery to the negatively charged cell membranes. [18]    The UV-vis-NIR spectra of CuS NPs, and PEG-CuS@MSNs and PEG-CuS@MSNs-DOX all showed strong absorbance at the NIR region between 800 to 1400 nm, which is within photothermal treatment range ( Figure 4). However, a closer look at the spectra of the silica nanoparticles revealed that there was a slight shift in the spectra, which might due to the change in the local refractive index of the surrounding medium. [19] In addition, PEG-CuS@MSNs-DOX also exhibited the The drug release behavior of PEG-CuS@MSNs-DOX was studied in the pH of 7.4, 6.0, and 5.0 buffer solutions over a 24-h period ( Figure 5). It can be seen that DOX release was pH dependent. At pH 5, the 24 h cumulative release of DOX for PEG-CuS@MSNs-DOX was 17.9%. At pH 6.0, the released ratio decreased to 12.4%, and dropped down to 5.7% when pH increased to 7.4. In addition, the release results also indicated that NIR laser irradiation accelerated DOX release in various pH conditions. Each time upon laser irradiation for 2 min at 2 W/cm 2 , released DOX is    The therapeutic efficacy test in vitro was explored on A549 cells incubated in 24-well plate. It was observed efficient photothermal ablation of the A549 cells only after 1 min irradiation of the 900 nm laser in the presence of the nanoparticles. Also, in the trypan blue assay as shown in Figure 7, few cells were dead either after laser exposure alone (7b) or after treated with different nanoparticles without laser exposure (Figure 7c and 7e). However, almost all the cells were dead after laser irradiation in all nanoparticle groups (Figure 7d 7f and 7g). This was attributed to the efficient intracellular uptake of PEG-CuS@MSNs-DOX after 2 h incubation.

CONCLUSIONS
In summary, monodisperse CuS nanocrystals coated in uniform pore-sized mesoporous silica nano spheres with an average particle size of 40 nm were successfully synthesized. Mesoporous silica spheres adsorbed doxorubicin and enabled high drug loading capacity. The release rate of doxorubicin was faithfully controlled by pH, laser exposure and the surface property of mesoporous silica. They showed photothermal effects on cancer cells upon laser exposure. These mesoporous silica nanoparticles provide a facile pathway for versatile biomedical applications.

Introduction
The development of transdermal drug delivery systems (TDDS) is attractive because skin is the largest organ. TDDS have the distinct advantages over oral administration or injections since they directly deliver drugs into the skin or even the systemic circulation, avoiding first-pass clearance of liver thus enhancing bioavailability. TDDS provide sustained and steady-state pharmacokinetics, therefore decreasing administration frequency and improving the patient compliance. Further, TDDS avoid the limitation of injections such as pain, accidental needle-sticks, and possible side effects due to transiently high plasma drug concentration [1][2][3].
However, the skin presents a natural barrier to protect our body from the rough environment. It forms multilayers in the epidermis, which include stratum corneum (SC), stratum lucidum, stratum granulosum, stratum pinosum, stratum basale from topical toward dermis. The SC is the outmost layer and consists of dead keratinocytes or corneocytes intercalated with lipids [4]. This 10-to 20-µm thick layer is the formidable barrier preventing most drug molecules from permeation. Only lipophilic drug with molecular weight (MW) less than 500 Daltons is able to penetrate the skin barrier, such as clonidine, fentanyl, and lidocaine [3,5].
A variety of methods have been tried to enhance the permeability of the SC. Chemical enhancers promote the drug penetration through the SC by disrupting the highly ordered bilayer structures of the intracellular lipids in the SC [3]. Conventional chemical enhancers such as Azone (1-dodecylazacycloheptan-2-one) as well as newly developed biochemical enhancers like peptides are of interest [6,7]. However, chemical enhancement has been shown little impact on delivery of hydrophilic drugs and macromolecules and irritation to living cells in the deeper skin [3]. On the other hand, physical enhancement techniques including mechanical and thermal approaches have been used to make micrometer dimensions of disruptions to SC structures. These micro-scale disruptions create channels of sufficient dimensions for passage of macromolecules. The thermal ablation activated by microheaters, [8] radio-frequency [9][10][11], superheated steam ejectate [12] or laser [13][14][15][16] is non-invasive technique to selectively remove small portions of the SC. These perforations are temporary, since the layers of the SC are continually replaced through the natural process of desquamation [8]. Some of the physical enhancement technologies have been applied in clinical trials for TDDS such as BA058 transdermal microneedle patch [17], transdermal basal insulin patch with microporation [18], teriparatide acetate TDDS transdermal [19], and electroacupuncture for opioid detoxification [20].
Laser ablation enhancement belongs to a physical approach that utilizes laser to perforate or remove the SC barrier in order to enhance the drug penetration. Water and pigments in the skin absorb the laser light energy and transform it into heat to achieve theromolysis of the skin. The heating duration must be controlled within microseconds in order to avoid heat propagation to deeper tissues [21]. The laser ablation approach enables precise control of depth of skin permeation, having the potentials for percutaneous delivery of biomacromolecules such as peptides, proteins, vaccines, DNAs [15]. In this review, we will focus on recent progresses of laser ablation enhanced TDDS.

Direct laser ablation enhancement
Although many types of laser with a broad wavelength range (193 -10,600 nm) are available in clinical practice such as ruby laser, neodymium:yttrium-aluminum-garnet (Nd:YAG) laser, alexandrite laser, CO 2 laser and erbium:yttrium-aluminum-garnet (Er:YAG) laser (Table 1), only a few are applied to transdermal delivery so far.
Pulsed CO 2 and Er:YAG laser are in common use for SC ablation [22]. The ruby laser (694 nm) and the alexandrite laser (755 nm) belong to near-infrared (NIR) laser (650 -900 nm). The NIR light causes little tissue absorption or minimal thermal effect [23], which is not sufficient to remove the SC. By contrast, the wavelengths of the CO 2 and Er:YAG laser are 10,600 nm and 2,940 nm, respectively. Both lasers directly induce heating and microporation of the skin through water excitation and explosive evaporation from the epidermis. Between these two laser types, the wavelength of the mid-infrared Er:YAG light matches a principal absorption wavelength for water molecules [13]. Compared with the CO 2 laser, the Er:YAG laser is about 15 times better absorbed in skin [22]. Therefore, the Er:YAG laser has a much higher ablation efficacy and a lower ablation threshold [24]. The Er:YAG laser shows the reduced thermal damage even in deeper crater holes in comparison with the pulsed CO 2 laser [22,24]. These favorable properties make the Er:YAG laser an ideal light source not only for skin surgery but also for enhanced transdermal drug delivery. A comparison of three sources of laser, the ruby, CO 2 and Er:YAG laser, on the skin permeability for 5-Fluorouracil (5-FU) showed that the ruby laser only moderately enhanced the drug flux [25]. The Er:YAG laser with fluence at 0.8 -1.4 J/cm 2 enhanced the flux of 5-FU by 53 -133 times than untreated skin. The CO 2 laser increased penetration of 5-FU by 36 -41 times under the fluences of 4.0 and 7.0 J/cm 2 with certain thermal effects [25].
Laser-induced thermal ablation heats the skin to hundreds of degrees for very short periods of time (micro-to milli-seconds) to disrupt the SC [3]. The extent of structure alteration of the SC is proportional to the temperature locally elevated, i.e. (i) disordering of SC lipid structure by temperature between 100 o C and 150 o C, (ii) disruption of SC keratin network structure by temperature between 150 o C and 250 o C, and (iii) decomposition and vaporization of keratin to create micron-scale holes in the SC by temperature above 300 o C [21]. Correspondingly, skin permeability was increased from a few fold to three orders of magnitude [21]. For thermal ablation-enhanced TDDS, high energy of laser with pulse duration less than microseconds is required because it generates limited or negligible heat transfer to surrounding tissue [13][14][15][16]. The microsecond-pulsed laser steepens the temperature gradient across the SC. The skin surface is extremely hot but not the viable epidermis and deeper skin tissues [12]. This technique referred to as "cold ablation", thereby, largely eliminates side effects and vastly improves safety.
In physically enhanced TDDS, the controllable depth and wound area of skin perforation by the laser ablation should be well considered. Based on the clinical data from microneedle and thermal ablation-enhanced transdermal delivery, micron-scale defects in the SC are well tolerated by patients as long as no significant damage to living cells in the viable epidermis and dermis [3]. To solve this issue, a laser microporation technology called P.L.E.A.S.E. ® (Precise Laser Epidermal System; Pantec Biosolutions) has been developed by using a diode-pumped fractional Er:YAG laser (Fig. 1A) [14,15]. Instead of conventional Er:YAG in clinics that ablates a 7-mm spot on the skin, P.L.E.A.S.E. ® generates a matrix of identical micropores with 100 -150 µm wide of each (Fig. 1B). Since the concentrated laser beam are divided into microbeams, P.L.E.A.S.E. ® efficiently and fractionally ablates skin with less damages (Fig. 1C) [14]. In addition, the pulse duration of the fractional laser from P.L.E.A.S.E. ® is shorter than conventional Er:YAG laser to assure the localization of heat transfer to the skin surface without allowing heat to propagate to the viable tissues below. This technology is patient-friendly since it is programmed to precisely control the number of micropores in unit area and depth of micropores based on the laser fluence [15].

Photothermal nanoparticle-mediated laser ablation enhancement
The development of nanotechnology brings a breakthrough to the limited application of NIR laser in TDDS. Gold nanostructures such as nanoshells [26], nanorods [27], nanocages [28,29], and hollow nanospheres [30] possess unique optical properties due to strong and tunable surface plasmon resonance (SPR). They can be synthesized to specifically absorb NIR light and convert photo energy into thermal energy to raise the temperature of surrounding tissue [26,31]. Nanoparticles with the property of photothermal coupling effect are called photothermal nanoparticles. Gold photothermal nanoparticles can be applied to photothermal ablation therapy of tumor cells [32][33][34][35], as well as the NIR laser-controlled drug release [36][37][38][39][40]. The absorbance of NIR light is desirable because it causes minimal thermal injury to normal tissues with optimal light penetration [23,41]. Recently, a surfactant/protein/gold nanorod complex has been applied to transdermal delivery of proteins [42]. The solid-in-oil dispersion system has been formulated through incorporation of gold nanorods as the photothermal ablation enhancer to disrupt the skin barrier. This approach effectively enhances the protein permeation through the skin in vitro and induces an immune response in vivo [42]. In this application, instead of pulsed laser, a xenon lamp that required high light power (6 W/cm 2 ) and long duration of light exposure (20 min) has been used to ablate the stratum corneum [42]. Therefore, the heat propagation to the deeper tissue could be a major concern.
Semiconductor CuS nanoparticles (CuSNPs) are a new class of photothermal nanoparticles that provide an alternative to gold analogs. Compared to gold, CuS is much less expensive [43]. Irradiated with NIR laser, CuSNPs generate heat for photothermal destruction of tumor cells [43][44][45][46]. Hollow CuSNPs (HCuSNPs) have been utilized for photothermal ablation-enhanced transdermal drug delivery [47]. A nanosecond-pulsed Nd:YAG laser in tandem with Ti:Sapphire laser (900 nm) has been used to induce rapid heating of the nanoparticles and instantaneous heat conduction.
Such type of laser with nanosecond pulse duration provides focused thermal ablation of the SC and minimizes skin heat accumulation. The average temperature of the irradiated skin area only increases to ~40 -50 o C. The depth of skin perforation can be precisely controlled by adjusting the laser power. The skin disruption by HCuSNPs-mediated photothermal ablation significantly increases the permeability of macromolecule drugs, providing effective percutaneous delivery [47].
Dextran, a hydrophilic macromolecular model drug, was used to evaluate the skin permeation. By using a laser with fluence above 1.7 J/cm 2 , the transdermal transport of FDs with molecular weight ranging from 4.4 kDa to 77 kDa was significantly enhanced. The possible mechanism could be ablation of the SC layer, photomechanical stress on intercellular regions, and alterations of the morphology and arrangement of corneocytes by the Er:YAG laser. Further, the transdermal delivery of hexameric insulin was higher than that of 38-kDa FD, suggesting the potential of laser ablative transdermal delivery of Insulin [13].
ATG and Basiliximab, two marketed antibodies for the induction of immunosuppression, were studied with fractional Er: YAG laser [15]. The result showed that the increase of pore numbers and laser fluence promoted the transdermal permeation of the antibodies. Total delivery of ATG at 24 h after laser treatment (900 pores, at a fluence of 45.3 J/cm 2 ) increased 82.8-fold over the control (untreated skin).
Increasing laser fluence from 22.65 to 135.9 J/cm 2 enhanced total ATG delivery from 1.70 ± 0.65 to 8.70 ± 1.55 µg/cm 2 , respectively. Similar penetration enhancement was observed in Basiliximab. Moreover, the in vitro and in vivo result was well correlated in a mouse model [15].
Topical delivery of DNA and RNA were also enhanced by laser ablation [53,55].
With Er:YAG treatment, in vitro permeation of antisense oligonucleotides (ASOs) increased 3 -30-folds, depending on the laser fluence and the molecular weight of ASO. In vivo results showed an enhanced expression of plasmid DNA in the epidermis and subcutis [53]. Besides, it was also found that the delivery rate of siRNA was raised by several times by the laser application [55].
Laser-enhanced transcutaneous protein delivery provided a non-invasive immunization method [15,56,57]. The laser induced microporation allowed high levels of antigen uptake. Further, transdermal delivery of vaccine targets the potent epidermal Langerhans and dermal dendritic cells that generate a strong immune response at much lower doses than hypodermic injection [58]. Transcutaneous application of OVA via laser-generated micropores using the P.L.E.A.S.E ® device induced equal or higher immune responses compared to immunization by s.c. injection [57]. In addition, targeting different layers of the skin had the potential to bias different T cell polarization patterns [57]. The laser ablation enhancement followed by transcutaneous immunization of lysozyme with 129 amino acids (14,307 Da) induced antigen-specific IgG in the serum by 3-fold compared to the control without laser treatment [56].
In addition to deliver drug compounds, laser ablation-enhanced transdermal delivery of adipose-derived stem cells (ADSC) were explored for wound healings [54]. After fractional Er:YAG laser treatment, bromodeoxyuridine (BrdU)-labeled ADSC was applied to the laser treated areas. After 4 and 48 hours, 12% and 5.5% of the stem cells were found in the pretreated tissue, respectively [54]. This encouraging result furthered the studies to optimize the technology for future clinical applications.
Because of high photothermal conversion effect, the gold nanoparticles were utilized to achieve thermal ablation of skin to enhance transdermal delivery of OVA [42]. In this study, a solid-in-oil dispersion was formulated to incorporate both the gold nanorods and the drug. Therefore, the nanodispersion exerted two modules upon NIR light irradiation, i.e. thermal ablation of the SC by the gold nanorods and enhancement of skin permeation of OVA. In vivo experiment showed significant increase of immune response for the gold nanorod-OVA solid-in-oil dispersion with NIR light treatment than other groups [42]. Another study investigated the use of HCuSNPs as photothermal ablation enhancers [47]. The permeability of human growth hormone (hGH) in skin applied with HCuSNPs plus NIR laser was increased by 3 orders of magnitude in comparison with that of the intact skin. In vivo study showed that transdermal delivery of hGH using the HCuSNP-mediated photothermal ablation technique reached an average bioavailability of 83% relative to that of the subcutaneous injection. The peak drug concentration through transdermal delivery was only one-third of that via subcutaneous delivery [47]. This was clinical beneficial because it reduced the risk of side effect related to high concentrations and controlled the drug concentration in a relatively stable level.

Conclusion
In conclusion, this review has discussed recent progresses of laser ablation technology to enhance transdermal drug delivery. The success of delivery relies on locally thermal ablation of the SC. By adjusting the laser fluence and exposure time, the depth of the microporation can be controlled without harming the deeper living tissues such as the dermis. The microchannels allow skin permeation of hydrophilic and macromolecular compounds. Particular interest has been shown in the development of the photothermal nanoparticles that mediate photothermal ablation of skin and deliver drug in a single setting. As a clean, needle-free and non-invasive approach, laser ablation enhancement technology shows great potential for future market.    less the damage to healthy cells due to the precise delivery of high-energy radiation to one particular location. Radiation nevertheless presents short-term risks such as skin rash or problems with tissues or organs near radiation pass and long-term side effects such as infertility or even secondary cancer due to radiation exposure.
The heat generated by laser has been used frequently in clinic practice. For example, in dentistry, laser is used for photobiomodulation in treatment of recurrent aphthous stomatitis and traumatic ulcers to accelerate healing process [3]. It is also used in thermal therapy to heat malignant tissue and tumor. Other examples include laser coagulation to seal blood vessels and stop bleeding, laser welding to join tissues and blood vessels, laser shock waves to remove urinary, kidney and biliary stones [4].
The detailed strategy and outcome of PTT on tumor treatment is largely depending on the properties of tumor, such as location, size, surface characteristic, water content, etc. Generally, laser can only reach less than a few (3-4) centimeters under skin [5].
For superficial tumors, external heating could be achieved by superficial applicators.
For tumors deep inside the body but not close to a body cavity, a fiber need be inserted into the center of the target tumor [6]. Moreover, it is important to keep well-localized heating high enough in the tumor but not harm the surrounding normal tissue. To achieve this, additional use of real time imaging techniques such as MRI and ultrasound imaging are required.
PTT has been applied to various solid tumors, including penile cancer, bladder tumors, renal tumor, melanoma, cervical cancer, breast cancer and so on. For selected cases of penile primary cancer in early stage, Nd:YAG or CO 2 lasers therapy was found effective [7,8]. For bladder and renal cancer, lasers is feasible for resection, coagulation, and enucleation of non-muscle invasive bladder tumors, but it should only be used in clinical trial setting or for patients who are not applicable to conventional treatment either because of co-morbidities or other complications [9].
For cutaneous melanoma, multiple small (<1 cm) lesions respond well to the CO 2 laser. CO 2 laser is recommended or palliative treatment of locoregional recurrence in a limb [10]. Local laser ablation can also be used on precancerous dysplasia to prevent cervical cancers [11].
There are also a lot of clinical trials for PTT on various carcinomas. For brain tumors, patients with recurrence of glioblastoma who had previous received total resection, chemotherapy and radiation therapy and ineligible for a secondary surgery was proposed for MIR guided laser-induced thermal therapy (LITT) salvage therapy [12]. Patients with resistant metastatic intracranial tumors who had previously undergone chemotherapy, whole-brain radiation therapy, and radiosurgery were given real-time magnetic resonance-guided laser-induced thermal therapy. The procedure was safely carried out with minimal invasion in one day [13]. Other MRI-guided laser interstitial thermal therapy was investigated on liver metastasis and prostate cancer [14][15][16]. A minimal invasion method, percutaneous laser ablation, by inserting optical fibers into the cancer through 21-gauge needles, is safe and effective for cirrhotic patients with hepatocellular carcinoma when resection or liver transplantation is not possible [17]. In addition, tumor caused thrombus and mucostis are also eligible for laser therapy [18,19].
Normally, tumor larger than 3 cm is not proper for laser ablation. However, a drug might change this situation. Sorafenib, a multikinase inhibitor to reduce intratumoral blood flow, might enhance the effectiveness of laser ablation on hepatocellular carcinoma larger than 4 cm by decreasing cancer microvessel density and thus enlarge laser-induced coagulation necrosis area [20].
Recently PTT has attracted new interest because of the arising of photothermal nanoparticles, especially gold colloidal. Gold nanoparticles have superior light absorption the excited electrons on the surface can produce strong localized heat thus damage the cancer cell. Gold nanoparticles have involved in clinical trials. The first example, which has phase II result, is gold nanoparticles with silica-iron oxide shell for PTT treatment of atherosclerosis. The gold nanoparticles were integrated in stem cells grown on a bioengineered patch. Then the patch was implanted onto the artery through the minimally invasive cardiac surgery. Under near-infrared laser irradiation, the nanoparticles in the patch can burn the plaque. The dense calcium area, fibrous and fibro-fatty tissue with fulminant necrosis significantly decreased due to thermolysis of the nanoparticles after 12 months [21]. In addition, Tumor necrosis factor (TNF)bound colloidal gold (Aurimmune®) was under phase I clinical trial. Patients with advanced solid organ malignancies or primary and metastatic cancer undergoing surgical resection received colloidal gold-bound TNF intravenously 12-78 hours prior to surgery. The antitumor effect and biodistribution is still under evaluation [22,23].
Another gold nanoparticle based laser photothermal therapy (Aurolase®) is also under clinical trial for patients with refractory and/or recurrent tumors of the head and neck and subjects with primary and/or metastatic lung tumors. The gold-silica (Auroshell®) nanoparticles accumulated at the tumor respond to the interstitial illuminations of an 808-nm laser [24,25].
Besides gold nanoparticles, other types of nanoparticles are also widely explored.
For example, graphene, two-dimensional (2D) crystal of sp 2 -hybridized carbon atom arranged in six-membered rings with high optical absorption in the NIR region are utilized for PTT to ablating tumor [26,27]. Semiconductor nanocrystals such as CdS, CdSe, CdTe, have been intensively used for fluorescence bioimaging due to the size and shape-dependent quantum confinement effect [28]. CuTe, etc [29,30]. Inorganic nanoparticles is hard to degrade in vivo and might possess potential long-term toxicity. As a promising substitute, organic nanostructures are explored for photothermal therapy based on the discovery of NIR-absorbing organic nanomaterials. For instance, indocyanine green (ICG) has been approved by FDA. Other NIR dyes include heptamethine indocyanine dye-IR780, IR783, IR808, IR825, PcBu4, porphyrins [31][32][33][34][35][36]. In order to prepare more stable formulation with various sophisticated purpose, the NIR dyes incorporated into micelles, liposomes or even proteins, have been used for photothermal tumor ablation [37][38][39]. Conjugated polymers such as polyaniline, polypyrrole and PEDOT: PSS-PEG with extended π-electrons also show high NIR absorbance as PTT agents, and have been found to be robust photothermal agents [40][41][42]. In addition, metallic nanoparticles, mesoporous silica nanoparticles and rare earth doped nanoparticles were also explored for photothermal therapy of cancer [43,44]. It is not possible to highlight just one nanoparticle since their parameters are so different. However, they should meet the general requirement for in vivo application-deep tissue penetration ability for the corresponding wavelength, significant photothermal conversion efficiency and decent biocompatibility.
PTT is effective for recurrent tumors not sensitive to chemo drug or radiation.
However, PTT may not be effective against all -particularly cancers that have already metastasized all over the body. Chemo drugs, in contrast, are able to distribute throughout the body to destroy cancer cells that have spread, although chemo drugs might cause huge side effects. The combination of PTT and chemotherapy might conquer the limitation of PTT and chemotherapy. Doxorubicin (DOX) is a chemo drug used for various cancers, such as blood cancers, many solid tumors and soft tissue sarcomas. It works by intercalating DNA and increasing free radical. DOX has a 'lifetime maximum dose' due to its life-threatening dosage-dependent cardiac toxicity [45]. In addition to DOX, other chemo drugs were tested for synergistic enhanced phothermal therapy effect. For instance, small molecule chemo drugs like camptothecin, a DNA enzyme topoisomerase I inhibitor, was loaded into hollow copper sulfide nanoparticles for synergistic PTT therapy of cancer cells in vivo [46].
Hydrophobic curcumin (regulation multiple cell signaling pathways of tumor cells) was combined in Ag/Au nanogels for PTT [47]. Rotenone, a mitochondrial complex I inhibitor, [48] and cisplatin (binding to and causing crosslinking of DNA) and so far were also explored [49]. Besides, oligodeoxynucleotides containing cytosine-guanine (CpG) motifs (a small single strand DNA functions as immunoadjuvant), single and double strand DNA [50], tumor necrosis factor (TNF), antibody anti-VEGFR2, etc were all investigated [51].
The CuS nanoparticles are considered as potent counterpart to gold nanoparticles, since they are economically efficient, not affected by the surrounding (The PTT of CuS nanoparticles originated from d-d* electron transition instead of surface plasmon resonance effect), low long-term toxicity [52]. Various Cu 2-x S nano composites (e.g. chelator-free [(64)Cu]CuS nanoparticles, hollow nanoshell, core-shell nanoparticle [46,53,54] have been extensively explored as a transformer of laser to thermal heat either accountable for therapeutic effects or as vehicles for drug delivery.