text:
Fabrication and characterization of a novel self-assembling micelle based on chitosan cross-linked pectin–doxorubicin conjugates macromolecular pro-drug for targeted cancer therapy

Cancer is one of the leading causes of morbidity and mortality worldwide. Doxorubicin is one of the most effective anticancer drugs approved by FDA. However, like all the other anticancer drugs, the efficacy of DOX is associated with high systemic toxicity to healthy tissues. In this study, chitosan cross-linked pectin–doxorubicin conjugates macromolecular pro-drug (CS-PDC-M) was prepared to enhance the therapeutic effects on liver cancer. CS-PDC-M was characterized in terms of size, size distribution, zeta potential, scanning electron microscope (SEM) and drug loading content. The CS-PDC-M achieved prolonged releasing ability was demonstrated by the in vitro drug release and in vitro cellular uptake assay. Biocompatibility of CS-PDC-M was screened by hemolysis activity examination, BSA adsorption test and cell viability evaluation in endothelial cells and LO2 cells. The CS-PDC-M achieved significantly high antitumor efficiency and targeting efficiency, which was demonstrated by the in vitro MTT assay and cellular targeting assay toward HepG2 cells, MCF-7 cells and A549 cells. The in vivo antitumor efficacy of CS-PDC-M was studied in athymic BALB/c nude mice bearing HepG2 cell xenografts. The organ damage assays of CS-PDC-M was studied in SD rats. Compared with that of free DOX and PDC-M, the CS-PDC-M exhibited higher antitumor efficacy and lower toxicity, implying that CS-PDC-M is a highly promising drug delivery system for hepatocellular carcinoma treatment.

Introduction
According to the United States Cancer Statistics Incidence and Mortality Web-based Report, cancer is a worldwide public health threat, and one of the leading causes of death. 1 To date, the conventional treatment on cancers is the removal of the tumor followed by extensive chemotherapy, radiotherapy, and immunotherapy. Chemotherapy remains the mainstay of treatment, and systemic chemotherapeutic agents include doxorubicin (DOX), 2-4 5-uorouracil, 5,6 paclitaxel, 7-9 and camptothecin. 10 DOX, an important anthracycline derivative 11 and a broad spectrum anticancer drug 12 approved by FDA, is widely used in oncology. [13][14][15] Intracellularly, DOX intercalates into the DNA duplex and disrupts the replication and transcription process in cancer cells, leading to cell apoptosis. 16,17 Although highly effective in cancer therapy, DOX is strictly limited in its clinical use due to its poor target selectivity, and serious side effects such as cardiotoxicity and cytotoxicity to normal tissues. 18,19 Various polymer-based drug delivery systems with good biocompatibility and biodegradability have been developed in order to reduce side effects and improve the therapeutic. 3,16 There are several types of polymer-based drug delivery systems that could be exploited for the relevant purposes, including polymer-drug conjugates, [20][21][22] micelles, 23,24 and vesicles. [25][26][27] Pectin-doxorubicin conjugated macromolecule prodrug (PDC-M) nanoparticles, a novel amphiphilic core-shell micelle, was successfully synthesized. The PDC-M demonstrated a longer half-life, better hydrophilicity and lower toxicity compared to DOX. 28 The surface charge of PDC-M is negative because the pectin contained multiple negatively charged saccharides carboxylic. 29 The incorporation of positive charge onto the surface of polymer materials can effectively improve the cell attachment capacity of the polymer materials without a negative effect on cell viability instead. Moreover, the adhesion of cell properties can be signicantly promoted with the positive charge modications on the surface of polymer-based drug delivery systems. 30,31 It is necessary to nd a material to modify the surface of PDC-M.
Chitosan (CS), generally used as biomaterial in pharmaceuticals, is a natural positively charged polysaccharide. 32,33 Under acidic condition, the amino groups of CS are protonated and become positively charged species. 34 In addition, CS is widely used in pharmaceutical and biomedical applications because of its favorable properties: biocompatibility, biodegradability, non-toxicity, antibacterial activity 35,36 and lowimmunogenicity. 37 In this work, CS-PDC-M was synthesized and prepared through incorporating the positively charged chitosan onto the PDC-M surface with negatively charged pectin. The physicochemical properties and characteristics of CS-PDC-M, including morphology, particle size, zeta potential, in vitro drug release were investigated in this paper. Biocompatibility of CS-PDC-M was screened by hemolysis activity examination, BSA adsorption test and cell viability evaluation in endothelial cells and LO2 cells. The in vitro antitumor effects of CS-PDC-M were evaluated by the MTT assay, cellular targeting assay toward HepG2 cells, MCF-7 cells and A549 cells. Cellular uptake was evaluated in HepG2 cells. Finally, in vivo tumor growth inhibition was analyzed in tumor-bearing athymic BALB/c nude mice bearing HepG2 cell xenogras. The organ damage assays of CS-PDC-M was studied in SD rats.

Materials
Chitosan (CS, Biomedical Grade, M w ¼ 5 Â 10 5 , degree of deacetylation is 91%) were supplied by Golden-Shell Pharmaceutical Co., LTD, Zhejiang, China. Low methoxyl pectin (pectin from citrus peel, esterication degree less than 50%, galacturonic acid 75.0%, M w ¼ 1.18 Â 10 5 ) and doxorubicin (DOX) were obtained from the Milky Way Chemical Co., Ltd. Hubei. Dulbecco's modied Eagle's medium (DMEM) and Fetal Bovine Serum (FBS) were obtained from Gibco (America). [3-4,5-Dimethylthiazol-2-yl]-2,5-diphenyltetrazolium bromide (MTT) was purchased from Amresco (Solon, OH). Human Umbilical Vein Endothelial Cells (HUVEC), LO2 cells, HepG2 cells, MCF-7 cells and A549 cells were obtained from Shanghai Life Sciences Academe, Shanghai, China. BALB/c nude mice were obtained from Beijing Vital River Laboratory Animal Technology Co., Ltd. Sprague-Dawley (SD) rats were supplied by Laboratory Animal Center, University of South China, China. All animals were kept in a pathogen-free environment and fed ad lib. The procedures for care and use of animals were approved by the Ethics Committee of the University of South China and all applicable institutional and governmental regulations concerning the ethical use of animals were followed. Human fresh blood is supplied by the Hengyang Hospital of Traditional Chinese Medicine. The research adhered to the tenets of The Declaration of Helsinki. The Ethics Committee of Hengyang Hospital of Traditional Chinese approved this study. Informed consent was obtained from all subjects before study. All other reagents were of analytical grade. All solutions were freshly prepared using Millipore-Q water, Billerica, MA, USA.

Fabrication of chitosan cross-linked PDC-M
According to our previous research, 28 the synthesis of PDC macromolecular pro-drug has been described by using the condensation reaction of the amide. Briey, 50 mL lowmethoxylated pectin solution (2 mg mL À1 ; pH 5.0) and 50 mL DOX solution (1 mg mL À1 ) was mixed. Then a mixture of EDC (100 mg) and NHS (30 mg) was added and maintained for 8 h at 50 C under magnetic stirring at 600 rpm. The core-shell PDC was fabricated during this process. Then the obtained PDC were washed with dehydrated alcohol and dialyzed in water for puried. To cross-link the PDC-M with chitosan, 1 mL chitosan solution (1 mg mL À1 ) was added dropwise to 10 mL PDC, at an average rate of 2 mL min À1 , using a nozzle with a 0.8 mm inner diameter. The CS-PDC-M was fabricated with magnetic stirred for 8 h. The resulting viscous suspension was centrifuged at 4000 rpm for 10 min to separate the chitosan-coated nanoparticles. Then the obtained CS-PDC-M was lyophilized. In vitro drug release study

Drug loading content and encapsulation efficiency
The drug release from the CS-PDC-M in vitro was evaluated using the dialysis method. During the process, CS-PDC-M (2 mL) was put into a dialysis bag. Then, the dialysis tubes were immersed into 30 mL of aqueous medias with different pH values (Tris-HCl buffer with a pH of 5.0, 6.8 and 7.4), and stirred at 200 rpm at 37 C for 240 h. At the predetermined time points, the entire buffer solutions in the outer tube were drawn out and replaced with 30 mL of fresh buffer solution. The DOX concentration in the removed buffer solution was subsequently detected using a UV-vis spectrophotometer by recording the optical density at 480 nm. Finally, the cumulative percentage of drug release at a specic time interval was calculated on the basis of the total amount of DOX loading measured at the beginning of this study.

Hemolysis activity examinations
Human fresh blood was rst centrifuged at 5000 rpm for 10 min to remove the supernatant and then was successively rinsed with 0.9% NaCl and diluted to 2% (v/v). Different concentrations of PDC-M and CS-PDC-M (19-304 mg mL À1 ) which have been adjusted to isotonic were mixed with 0.5 mL of 2% red blood cell suspensions. The deionized water was used as a positive control and 0.9% NaCl was used as a negative one. The mixtures were incubated at 37 C for 4 h and then were centrifuged at 3000 rpm for 5 min. The absorbance of the supernatants was recorded using a microplate reader (Labsystems Multiskan, Biochromatic Labsystem, Osaka, Japan) at 540 nm. The hemolysis ratio (HR) was calculated according to the eqn (3): In eqn (3), A sample , A negative control and A positive control were denoted as the absorbance of CS-PDC-M, negative and positive controls, respectively.

Bovine serum albumin (BSA) adsorption
To evaluate the adsorption of bovine serum albumin (BSA), 0.5 mL of PDC-M, CS-PDC-M and free DOX were mixed with 5 mL of BSA (0.5 mg mL À1 ) at room temperature. At specied time intervals, a small portion of this solution was removed and measured using the Bradford assay by UV-vis spectrophotometer at a wavelength of 595 nm. The amount of BSA adsorbed in samples was calculated from a standard calibration curve. All experiments were performed in triplicate at room temperature.

Cell viability assay
The in vitro cytotoxicity of PDC-M and CS-PDC-M were assessed using MTT assays in human umbilical vein endothelial cells, LO2 cells, HepG2 cells, MCF-7 cells and A549 cells. Typically, adherent cells (6 Â 10 3 cells per well) were evenly planted into 96-well plates and incubated overnight. Then cells were exposed to serial dilutions of free DOX, PDC-M and CS-PDC-M followed by 24 h incubation. Thereaer the plates were centrifuged to collect the cells, which were then washed three times with phosphate-buffered saline (PBS) and incubated for another 4 h with fresh medium containing 5 mg mL À1 MTT. Finally, dimethyl sulfoxide (DMSO) was used to dissolve the formazan crystals. The absorbance was determined at 570 nm with a microplate reader (Labsystems Multiskan, Bio-chromatic Labsystem, Osaka, Japan). Cell viability was calculated according to the eqn (4): In eqn (4), A sample and A control represent the absorbance of sample and control wells, respectively.

Cellular uptake
The cellular uptake was investigated by treating the HepG2 cells, MCF-7 cells and A549 cells with free DOX, PDC-M and CS-PDC-M. The cells were seeded in 6-well plates at 2.0 Â 10 5 cells per well in 2.0 mL of complete Dulbecco's modied Eagle's medium (DMEM) and cultured for 24 h, and then incubated at 37 C for additional 2 h with free DOX, PDC-M and CS-PDC-M in complete DMEM in the absence or presence of galactose. Then, the culture medium was removed and cells were washed with PBS thrice. Thereaer, the cells were xed with 4% (w/v) paraformaldehyde for 10 min at 25 C, and the cells were then washed three times in PBS and stained with DAPI for another 10 min. Aer that, the cells were washed with cold PBS three times to remove free DAPI and visualized under uorescence microscope (Olympus, Tokyo, Japan) with the following channels: red channel (DOX) excited at 488 nm, blue channel (DAPI) excited at 358 nm.
Quantitative analysis was performed by ow cytometry. HepG2 cells were cultured in six-well plates at a density of 1 Â   cells were trypsinized and collected by centrifugation as a cell pellet. Finally, the cells were suspended in PBS and observed by a ow cytometer (Beckman, California, USA) at an excitation wavelength of 488 nm and emission wavelength of 590 nm.

In vivo antitumor efficacy
The antitumor efficacy of free DOX, PDC-M and CS-PDC-M were evaluated in athymic BALB/c nude mice inoculated subcutaneously in the ank of the right fore leg with 0.2 mL of PBS containing 2 Â 10 6 HepG2 cells. The tumor volumes were monitored at predetermined intervals using a digital caliper. When the tumors grew to approximately 200 mm 3 (15 days aer tumor implantation), the mice were randomly divided into four groups (5 mice per group): saline (control), free DOX, PDC-M and CS-PDC-M. The mice were injected intravenously once every 2 days via the tail vein with saline (control), free DOX, PDC-M and CS-PDC-M at 5 mg kg À1 of DOX equivalents. Antitumor activity was evaluated in terms of tumor volume, which was calculated using the following formula: where V is the tumor volume, A is the length of the major axis and B is the length of the minor axis. The body weight was measured simultaneously as an indicator of systemic toxicity. The antitumor activity was evaluated via tumor growth and terminal tumor weight. Heart and tumor were collected 23 days aer treatment. Tissue slices were subjected to hematoxylin and eosin (H&E) staining.

Organ damage assays
The organ damage assays were performed by evaluating corresponding functional enzymes with fully automatic clinical chemistry and immunoturbidimetric analyzer (Liasys, AMS Alliance). The SD rats were weighed and randomly divided into 4 groups (6 mice per group): saline (blank control), free DOX, PDC-M and CS-PDC-M. The tail-vein injections of free DOX and nanomedicines with equivalent DOX (2.4 mg kg À1 ) in 1 mL of PBS or 1 mL of PBS alone were performed on days 0, 3 and 5 respectively. The blood of each SD rats was collected 7 days aer the treatment. Aerwards, the blood was centrifuged at 4000 rpm for 10 min to obtain plasma for measuring clinical parameter measurement, including alanine aminotransferase (ALT), aspartate aminotransferase (AST), blood urea nitrogen (BUN), creatinine (Cr), creatine kinase (CK), creatine kinase-MB (CK-MB) and lactate dehydrogenase (LDH).

Statistical analysis
Statistical comparisons were performed with one way ANOVA. p < 0.05 was considered to be statistically signicant. All values were expressed as mean value AE standard error (S.E.) of three separated experiments. *p < 0.05 or less was considered statistically signicant.

Results and discussion
The preparation of CS-PDC-M As shown in Scheme 1, pectin and DOX were bound by amide condensation to form PDC macromolecular pro-drug. Given that chitosan exhibits a strong positive charges while pectin has negative charges, the PDC-M was cross-linked with chitosan through electrostatic interaction. The incorporation of chitosan onto the PDC-M could further improve the cell adhesion of the CS-PDC-M. Furthermore, the systems could be more stable aer the crosslinking with chitosan.

Characterization of CS-PDC-M
The CS-PDC-M was characterized in terms of size, size distribution, zeta potential, scanning electron microscope (SEM) and drug loading content. As shown in Fig. 1A and Table 1 Table 1.
The morphologies of CS-PDC-M were observed by SEM (Fig. 1B). The SEM images revealed that most CS-PDC-M were spherical in shape and relatively homogeneous in size distribution. The study assumes that ideal nanoparticles were able to accumulate in tumors through a leaky tumor vasculature by the enhanced permeability and retention effect (EPR), control drug release inside the tumors, and this will eventually improve therapeutic efficacy while reducing off-target toxicity. Therefore, these desirable properties indicated that CS-PDC-M could be a suitable candidate for the delivery of drugs.
The stability in biological milieu is of signicant importance when choosing an appropriate drug carrier. The particle size of CS-PDC-M was measured to evaluate the stability of nanomedicines under physiological pHs. As shown in Fig. 1C, the particle size of the nanomedicines did not show signicant change during the time of investigation. All the results have demonstrated that the CS-PDC-M and PDC-M have good stability in biological milieu and could form a stable suspension in biological uids with various pHs and tonicity. All these properties are highly desired for a drug nanocarrier.

In vitro drug release proles
The maximum absorption of free DOX at 480 nm was identied by the UV-vis absorption spectroscopy. This information was set as the foundation for the following estimation of the cumulative release of DOX. Fig. 1D shows an in vitro cumulative release    28 These results strongly indicated that the CS-PDC-M system processed a prolonged slow release effect. This release pattern was even lower than DOX release from PDC-M in our previous report. The rate decrease of drug release could be explained by the coating of carrier by chitosan.

Biocompatibility of CS-PDC-M
Hemolysis of the blood is the problem associated with bioincompatibility. Hemolysis rates (HR) of human fresh blood with PDC-M and CS-PDC-M were shown in Fig. 2A.  growth than that of PDC-M. This is probably because the incorporation of chitosan onto the PDC-M could further improve the cell adhesion of the CS-PDC-M.
We found that DOX displayed better in vitro antitumor efficacy than that of PDC-M and CS-PDC-M. However, free DOX can cause severe side effects such as cardiotoxicity and is susceptible to metabolism by the liver. In contrast, PDC-M and CS-PDC-M can effectively reduce the side effect of chemotherapeutic agents due to the sustained release of anticancer drugs and active targeting.

Cellular uptake
The extent of the internalization of free DOX, PDC-M and CS-PDC-M into HepG2 cells, MCF-7 cells and A549 cells was investigated qualitatively by uorescence microscopy. The ASGP-R is a recycling endocytotic receptor on the surface of mammalian hepatoma cells (e.g. HepG2 cells), which can recognize galactose-and N-acetylgalactosamine-terminaled glycoproteins. 41,42 To investigate the important role of galactose acid in the cellular uptake of PDC-M and CS-PDC-M, we conducted a receptor competition assay by selecting galactose acid as the competitive reagent. Consistent with our hypothesis, when PDC-M and galactose were added to the wells simultaneously, the uorescence intensity of HepG2 cells signicantly decreased. As shown in Fig. 4A and B, the presence of galactose signicantly inhibited the uptake of CS-PDC-M and PDC-M into HepG2 cells, while the presence of galactose almost did not affect the endocytosis of nanomedicines into MCF-7 cells and A549 cells. As a comparison, the presence of galactose never affected the cellular uptake of free DOX. It suggested that the cellular uptake of galactosylated nanomedicine was mediated by the ASGP-R. Quantitative analysis of ow cytometric ( Fig. 4C and D) indicated that CS-PDC-M exhibited a 1.3fold higher uptake by HepG2 cells than that of CS-PDC-M + galactose, conrming that galactose can enhance the endocytosis of nanomedicines through the specic interactions with HepG2 cancer cells.
The incorporation of positive charge onto the surface of polymer materials can effectively improve the cell attachment capacity of the polymer materials. To demonstrate the enhanced cell attachment capacity of CS-PDC-M to HepG2 cells, the cellular uptake and intracellular DOX release behaviors of nanomedicines were evaluated by uorescence microscopy toward HepG2 cells. It can be seen from Fig. 5A that a week intensity of uorescence could be detected aer incubation for 30 min, suggesting that the PDC-M and CS-PDC-M can enter The endocytosis of free DOX and CS-PDC-M was conrmed by ow cytometry. Fig. 6A and B showed the cellular uptake efficiency of the CS-PDC-M and the free DOX by HepG2 cells in 2 h, 4 h and 6 h. Quantitative analysis indicated that free DOX exhibited 2.58-, 1.83-, 1.23-fold higher uptake by HepG2 cells than that to CS-PDC-M aer the same time culture. These results strongly suggested that the CS-PDC-M had sustained and prolonged release, in consistence with the results from the in vitro drug release proles.

In vivo antitumor efficacy
The antitumor efficacy of PDC-M and CS-PDC-M was studied in athymic BALB/c nude mice bearing HepG2 cell xenogras. M. Although the tumor growth was inhibited by free DOX treatment, the body weight of this group dramatically decreased compared to other groups (Fig. 7B). It suggested that severe toxicity was induced by free DOX. Eventually, all the animals in the free DOX group in a dose of 5 mg kg À1 had to be terminated for humanitarian reasons on day 12. The terminated animals in the free DOX group in dose of 6 mg kg À1 at day 7 or 10 mg kg À1 at day 12 post-treatment due to serious systemic toxicity caused  by free DOX has also been shown by other researchers, such as Kiziltepe et al. 43 and Gou et al. 13 On the contrary, the treatment with PDC-M and CS-PDC-M appeared to be well-tolerated, the animals exhibited unobvious weight loss but higher survival rate compared to that of the control and free DOX groups. The result demonstrated the signicant decrease of the toxicity of DOX when DOX incorporated into the PDC-M and CS-PDC-M. Fig. 7D and E showed the tumor morphology and average weight of each group at the end of the experiment. The tumor size of the control group was obviously larger than that of the treated groups, which is a straight evidence to the tumorinhibiting effect of the PDC-M and CS-PDC-M. Compared to PDC-M, CS-PDC-M had higher in vivo safety and therapeutic activity in animal models.
To further evaluate the in vivo anticancer activity and toxicity of PDC-M and CS-PDC-M, histology analysis was also employed. The tumors and the hearts of the rats were collected and sectioned for hematoxylin and eosin (H&E) staining. As shown in Fig. 8A and B, treatments with PDC-M and CS-PDC-M for 12 days have not caused any signicant morphological changes in heart compared to the saline group. However, it is worth noting that the free DOX treated group (aer treatment for 12 days) had noticeable damage in both heart and tumor. Aer treatment with PDC-M and CS-PDC-M for 23 days, most of the tumor tissue cells were destroyed and became necrotic, whereas the heart retained unchanged for the control treated groups. Therefore, these results proposed that the PDC-M and CS-PDC-M were less susceptible to side effects and had the best antitumor effects on liver cancer.

Organ damage assays
In order to compare the systematic toxicities of free DOX, PDC-M and CS-PDC-M aer all treatments, alterations of clinical chemical parameters in SD rates aer drug treatment were determined, including ALT, AST, AST/ALT, CK-MB, CK, LDH, BUN and Cr. It is well known that these parameters are associated with the function or certain damage of liver (ALT, AST, AST/ALT), heart (CK-MB, CK and LDH) and kidney (BUN and Cr). As shown in Fig. 9, SD rates treated with free DOX showed increased ALT values (mean value 146.0 AE 20.59 U L À1 ) and CK-MB values (mean value 2644.4 AE 731.997 U L À1 ) in plasma, which is indicative of early signs of liver and heart damage. Besides, when the dose of DOX dropped from 5 mg kg À1 to 2.4 mg kg À1 , according to Tran et al. 44 mice still died before the end of the experiment in free DOX group owing to the toxicities of DOX. On the contrary, the parameters of PDC-M and CS-PDC-M groups substantially maintained at normal levels, indicating that both PDC-M and CS-PDC-M did not lead to serious damage in liver, kidney and heart of mice in the treatment duration.

Conclusions
In this study, we successfully synthesized a novel micelle based on the cross-link of chitosan and PDC macromolecular pro-drug (CS-PDC-M) for targeted liver cancer therapy. The DOX gave a sustained and controlled delivery in the CS-PDC-M. CS-PDC-M and PDC-M had a suitable size of 221.2 nm for passive targeting to tumor tissues by the EPR effect. It was clearly observed that CS-PDC-M is spherical shape via SEM. The in vitro drug release proles and cellular uptake assays indicated that CS-PDC-M had a prolonged releasing ability. The hemocompatibilities, BSA adsorption test, cell viability evaluation and tissue compatibilities of CS-PDC-M and PDC-M were revealed to be excellent, which rendered their potential for delivering bioactive substance in vivo. The in vitro cellular targeting assay demonstrated that CS-PDC-M and PDC-M managed to target HepG2 cells specically with higher efficacy. The in vitro cellular uptake assay indicated the incorporation of chitosan onto the PDC-M could further improve the cell adhesion of the CS-PDC-M. Both in vitro cell experiments and in vivo animal studies showed that CS-PDC-M exhibited higher antitumor efficiency compared to the PDC-M solution, while reduces systemic toxicity of DOX during the treatment. Therefore, CS-PDC-M is a potentially versatile drug delivery system for targeted liver cancer therapy. Next, we will construct the co-delivery system on the basis of CS-PDC-M for co-delivery of doxorubicin and miRNA-122 to treat cancer.

Conflicts of interest
The authors report no declarations of interest.


Quantiﬁcation of Two Fluorophores’ Concentration Ratio in a Mice Model in Preparation for a Proposed Method for Early Detection of Alzheimer’s Disease

: Many biomedical applications require concentration measurements of biological compounds, which may be achieved using targeted ﬂuorescent probes. It has been shown that the ratio between amyloid-Beta and tau protein in the cerebrospinal ﬂuid (CSF) is a good indicator of incipient Alzheimer’s disease (AD). We have previously proposed a method that can accurately estimate the concentration ratio of these two proteins without the need to collect CSF samples: Fluorescent probes are injected to the blood and bind to the CSF biomarkers. A miniature needle with an optical ﬁber excites the ﬂuorescent probes and collects the ﬂuorescence emission. The concentration ratio between the proteins is estimated, and used for diagnosis of incipient AD. We present here the results of the method’s concentration ratio estimation during trials performed on mice. Miniature tubes containing two ﬂuorescent probes in several concentration ratios were implanted in two locations in the mice: subcutaneously, and deeper in the abdomen. The ﬂuorescent probes were excited and the ﬂuorescence intensity was measured. The concentration ratios were estimated from the ﬂuorescence intensities with an average error of less than 20%. This study is a step towards the demonstration of the method in in vivo studies in mice with markers in the bloodstream/CSF.

Introduction
Alzheimer's disease (AD) is the most frequent type of dementia [1], which includes a variety of diseases such as AD, vascular dementia, dementia with Lewy bodies, Parkinson's disease, and more. Dementia is caused by the death or malfunction of nerve cells in the brain, which results in a decline in memory and in several cognitive abilities. In AD, the pathological abnormalities are characterized by a gradual loss of axons and neurons, together with the presence of neurofibrillary tangles and amyloid plaques in the brain. According to the new criteria for diagnosing AD [1], the disease includes three stages: preclinical AD, Mild Cognitive Impairment (MCI) due to AD, and dementia due to AD. During the preclinical AD stage, which may last up to 20 years, no cognitive decline and memory loss symptoms occur. However, the pathogenic process in this stage already induces measurable changes in the brain, in the cerebrospinal fluid (CSF), and in the blood. Therefore, diagnostic tools for early have described a high quantum yield NIR probe AOI987, penetrating the blood-brain barrier. They demonstrated the interaction of the probe with amyloid plaques in APP23 transgenic mice in vivo and showed increasing fluorescence intensity with increasing plaque load. Zhang et al. [25] designed and synthesized curcumin based NIR probes. The CRANAD-58 probe showed significant fluorescence property changes upon mixing with both soluble and insoluble Aβ species in vitro. They tested it in vivo and revealed that CRANAD-58 was capable of differentiating transgenic and wild-type mice as young as four months old. They have also designed the CRANAD-17 probe, which was capable of inhibiting Aβ42 cross-linking induced by copper. This raises a possibility for CRANAD-17 to be considered for AD therapy.
Okamura et al. [26] report the optical profiles of THK-265, which has high binding affinity to amyloid plaques. They demonstrated that amyloid deposits in AβPP transgenic mice could be detected after intravenous administration of THK-265. They reported significantly higher fluorescence intensity in the brain of AβPP transgenic mice than in those of wild-type mice. Schmidt et al. [27] have also used THK-265 to develop and test a procedure for cerebral amyloid deposits visualization. They demonstrated in vivo the correlation between the intensity of the THK-265 signal and the plaque burden, indicating its utility for direct monitoring of Aβ aggregation progression. Cui et al. [28] report the synthesis, characteristics, and biological evaluations of a series of donor-acceptor molecules (DANIRs) as smart NIRF imaging probes for Aβ plaques in AD. These probes were proved to have affinity to Aβ plaques ex vivo, in vitro, and in vivo, with a significant fluorescence intensity increase and a blue-shift in the emission spectra.
Other groups have tried to develop fluorescent probes with affinity to tau proteins. Maruyama et al. [29] developed a class of tau ligands (PBBs) for optical and PET imaging. These included PBB5, a near-infrared fluorescent probe, which could assess tau accumulation in living mouse models. They demonstrated strong fluorescence of PBB5 in Tg mice expressing a single human four-repeat tau isoform (PS19) relative to non-Tg WT mice in the central nervous system (CNS) suggesting a concentration of this tracer in the PS19 spinal cord. Kim et al. [30] designed a NIR ratiometric probe, CyDPA2 that targets Tau aggregates. They evaluated it in vitro and ex vivo and reported an enhancements of absorption ratio and fluorescence intensity in a p-Tau concentration-dependent manner. In addition, fluorescence microscopy and gel staining studies demonstrated CyDPA2-labeled Tau aggregates.
In a previous report [31], we proposed an optical method that enables measurements of the CSF biomarker concentrations utilizing fluorescence probes, without the need to perform a lumbar puncture. This method is less invasive and has minimal risk, pain, and side effects. The method includes several stages. First, fluorescent probes [23][24][25][26][27][28][29][30]32], which bind specifically to the two biomarkers in the CSF, are injected to the blood. Second, a miniature needle with an optical fiber is inserted into the epidural fat, which is located several millimeters from the CSF, without penetrating the dura matter. The fluorescent biomarkers are then excited by a laser, which reaches the CSF through the optical fiber located inside the needle. Finally, the fluorescence intensities of both biomarkers are measured and used for the estimation of the biomarkers concentration ratio.
Since this optical method does not require CSF collection, the dura matter is not penetrated and a smaller needle may be used. Therefore, the lumbar puncture risks, which include CSF leakage, traumatic taps, spinal cord injuries, and severe headaches, are significantly reduced. In addition, as opposed to other similar optical methods [12,13] the concentration ratio estimation using this method does not require a priori information about the tissue optical and anatomical properties.
In our previous report [31] we introduced a theoretical model that enables to create the relation between the measured fluorescence emission and the biomarkers concentration ratio. The method was validated using Monte Carlo simulations and its accuracy was tested in vitro using multi-layered tissue phantom. Here we further assess the method's accuracy using mice. Miniature tubes containing two fluorescent probes with several concentration ratios were implanted in the mice. The tubes were inserted into two different locations: subcutaneously, and deeper in the abdomen. The fluorescent probes were excited separately and their fluorescence intensity was measured. Using a simple calibration curve, the concentration ratio of the fluorescent probes in each tube was extracted. The extracted ratios were compared to the true ratios and an error analysis was performed.

Fluorescent Probes
Several fluorescent probes with high affinity to amyloid-beta and tau protein were described in the Introduction [23][24][25][26][27][28][29][30]32]. However, these probes are still under development and are not commercially available, off the shelf products, hence could not be obtained. Therefore, we decided to test our method and verify the accuracy of the concentration ratio estimation using two fluorescent probes that may be purchased easily.
The optimal fluorescent probes should have separate absorption spectra, both in the range of 750-850 nm, where the changes in the absorption and scattering properties of the tissue are negligible. According to this requirement, Cy7 azide and Cy7.5 azide (Lumiprobe, Hallandale Beach, FL, USA), with absorption peaks at 750 and 788 nm and emission peaks at 773 and 808 nm, respectively, were chosen. The extinction coefficients of Cy7 and Cy7.5 are 199,000 and 223,000 [M −1 cm −1 ] and the quantum yield of Cy7 is 0.3. The value of the quantum yield of Cy7.5 was not given by the manufacturer.
Fluorescent samples were prepared with different concentrations between 50 and 900 nM, mixed with concentration ratios (Cy7:Cy7.5) between 1:2 and 1:14. These ratios cover the range of expected concentration ratio between amyloid-beta and P-tau proteins, for MCI patients developing AD, other dementia, or stable MCI.

Animal Studies
In order to verify the method's accuracy in animal models, miniature tubes containing fluorescent samples were implanted subcutaneously and abdominally in mice. The miniature tubes were 7 mm long, with 1.7 mm outer diameter and 1 mm inner diameter. They were filled with approximately 5 µL of the fluorescent samples. The tubes were then sealed at both ends using adhesive.
Five C57BL/6 mice were anaesthetized using intraperitoneal injection of ketamine (100 mg/kg) and xylazine (10 mg/kg). Two incisions of 3-5 mm were performed on each mouse. One incision was made in the skin near the hind leg of the mouse. A subcutaneous pocket was created and a tube was inserted into the pocket. The pocket was then closed using a 9-mm AutoClip. A second incision was made in the abdomen area of the mouse. This incision was deeper than the subcutaneous one, enabling insertion of the tube a few millimeters deep in the mouse tissue, thus located below skin, muscle and fat tissues. This incision was also closed using a 9-mm AutoClip, which prevented any bleeding. The two fluorescent tubes measured directly after implantation, as will be detailed in the Sections 2.6, 2.7 and 3.2. Anesthetization lasted up to three hours, during which the implantation and fluorescence measurements were performed. When necessary, an additional dose of ketamine (50 mg/kg) was administered to the mouse subcutaneously. All mice were sacrificed by cervical dislocation. All the experimental procedures were approved by the Animal Care Committee of Tel Aviv University (approval number: 034_b8007_26).

Reference Phantom
A reference phantom required for the calibration process was composed of polyester resin with Titanium Dioxide, according to the protocol described by Firbank and Delpy [33]. Titanium dioxide was added in order to increase the scattering coefficient of the polyester, with a 0.53% titanium dioxide mass percentage. The reference phantom was squared with 2.7 mm thickness and the length of the side was 35 mm.

System Setup
The samples were excited with a mode-locked, Ti:Sapphire femtosecond laser (Tsunami, Newport/Spectra-Physics, Santa Clara, CA, USA), with 80 MHz repetition rate, pumped by a 5 W, 532 nm laser (Millennia vs., Newport/Spectra-Physics). The laser was tuned to 750 nm for the cy7 excitation and 800 nm for the cy7.5 excitation. The excitation beam was split to a fast photodiode (PHD-400, Becker & Hickl GmbH, Berlin, Germany) for signal synchronization, and to the center fiber bundle of a fluorescence inspection probe (Oriel 77558, Newport, Santa Clara, CA, USA). The Oriel inspection fiber has a central core with a fiber bundle that carries the excitation radiation to the samples. The core is surrounded by a ring of fibers that collect and deliver the fluorescence from the sample to the output channel. The source detector separation, between the central core (input) and the fiber ring (output), was less than 1 mm. The collected fluorescence is delivered through the distal end of the output channel into a filter wheel switching between a band-pass filter FF01-780/12 (Semrock, Rochester, NY, USA) for the Cy7 and long-pass filters RG830 and RG850 (Newport) for the Cy7.5 collection. The emission was measured using a fast photomultiplier tube (PMT) head (H7422P-50, Hamamatsu, Japan). The PMT was controlled using the DCC-100 detector controller (Becker & Hickl GmbH). Data acquisition was accomplished using the TCSPC module (SPC-730, Becker & Hickl GmbH) synchronizing between the SYNC signal arriving from the photodiode and the constant fraction discriminator (CFD) signal from the PMT.

Calibration Process
A calibration process was performed prior to the animal fluorescence measurements of the miniature tubes. For the calibration process five fluorescent samples were prepared, covering the range of ratios that were measured during the experiment. The samples were placed in sealed cuvettes and the reference phantom was placed above them. The fluorescence energy emitted by these samples was measured at two excitation wavelengths, one for each fluorophore, with the appropriate emission filter. For each measurement, the total number of photons measured by the TCSPC system was used as the fluorescence energy value. A calibration curve of fluorescence energy ratio versus concentration ratio was created based on these measurements.

Intensity Ratio Measurements
Five mice were anesthetized and implanted each with two fluorescent tubes-subcutaneous and deeper in the abdomen. Two of the tubes were excluded due to problems with the location of the implantation. Eventually the experiment included a total of eight tubes with eight different fluorescent ratios. The fluorescence emission of the probes in the tubes, covering ratios in the range of 1:2 to 1:14, was measured using the fluorescence inspection probe. Four tubes were located subcutaneously and four were located deeper in the abdomen of the mouse. When necessary, the mice were shaved prior to the fluorescence excitation.
The fluorescence measurements were repeated several times for each fluorescent tube. The location of the fiber bundle on the mouse skin surface was changed for each excitation in order to test varying tube depths. After each location change of the fiber bundle, a time-resolved fluorescence intensity was taken in order to ensure that the fluorescent tube is in the imaged photon path. In cases where very low fluorescence intensity was measured the fiber bundle was relocated. In other cases, high intensity was measured but a Gaussian-shaped, time-resolved curve without the expected fluorescence lifetime decay was measured, thus indicating that the collected photons are only resulting from laser reflections, and not from fluorescence. In those cases, the fiber bundle was relocated again until a typical fluorescent lifetime curve was measured. The measured energy ratio was then converted to a concentration ratio value using the calibration curve and compared to the true concentration ratio value.

Intensity Time-Resolved Measurements
Fluorescence intensity time-resolved measurements were taken from tubes implanted both subcutaneously and in the abdomen. The time-resolved curves were measured using the TCSPC system.

Concentration Ratio Estimation
Initially, a calibration process was performed using a polyester reference phantom before measuring the fluorescent samples in the mice. A calibration curve was created and used for estimating the concentration ratios of the fluorescent tubes located subcutaneously in the hind leg and deeper in the abdomen. Table 1 summarizes the data of the eight fluorescent tubes used in the animal experiments, detailing the number of the fluorescent tube, its location, its true and estimated concentration and the estimation errors. Figure 1 presents the estimated concentration ratio versus the true concentration ratio. Each data point is the average estimated value of four to six measurements (performed on different excitation locations on the mouse skin surface) with its 90% confidence interval. The total number of measurements performed on the eight tubes was 41. The data points of the tubes implanted subcutaneously are represented in Figure 1 using the star symbols and the abdomen tubes are represented using the circle symbols. The dotted lines represent the true ratio values with 15% error value and the dotted lines represent the true ratio values with 35% error value. The mean absolute error of the estimated values presented in Figure 1, is 19.86%, with a standard deviation of 14.04%. When calculating the mean error (as opposed to the mean absolute error), which averages the over estimations and the under estimations, the mean error is 0.69% with a standard deviation of 25.45%. This result indicates that there wasn't a consistent tendency neither to overestimation nor to underestimation. Therefore, we assume that the reference phantom had more or less the same attenuation as the mouse tissue, although it may have had a flatter absorption spectrum than the mouse tissue. Actually, if the results did show a clear trend for overestimation the accuracy could have been improved using a correction factor. Unfortunately, this cannot be done in our case, but additional experiments, with different reference phantoms, should be performed for complete validation.

Abdomen vs. Subcutaneous Tubes
The fluorescent tubes were implanted in two locations-subcutaneously and in the abdomen. It is interesting to investigate if and how these locations affected the accuracy of estimation. The tubes implanted subcutaneously were located in a depth of less than 1 mm under skin surface, whereas the tubes implanted in the abdomen were implanted in a depth of 1-5 mm. It is worth mentioning that in some cases where the mice fur was rather thick, the mice were shaved, while in other cases the fur wasn't removed. The shaving was performed in order to reduce the attenuation caused by the thick black fur. The shaving was performed both for the subcutaneous tubes and the abdomen tubes. Since the fluorescence intensity ratio was calculated, and it is assumed that the absorption and emission spectra of the fur is almost flat in the used wavelengths, the shaving should not affect the accuracy of the results. The shaving only assisted in achieving higher fluorescence intensities, reducing the measurement time and perhaps increasing the signal to noise ratio. In addition, although the distance of the subcutaneous tubes from the mouse skin was less than the distance of the abdomen tubes, the thickness of the fur in the subcutaneous tubes was higher than in the abdomen. This is due to the fact that the fur in the leg area, which was chosen for the subcutaneous implantation, is thicker than the fur in the abdomen area. Therefore, although the energy attenuation due to the mouse tissue was rather small in the subcutaneous tubes, there was additional attenuation due to the fur. When comparing the mean error in the subcutaneous tube with the abdomen tubes, as presented in Figure 2a, we notice a slight overestimation in the abdomen tubes (4.95% ± 31.22%) compared to a slight under-estimation in subcutaneous tubes (−3.58% ± 22.08%). However, in our opinion this is not due to higher effective attenuation in the subcutaneous tubes, since in most subcutaneous tubes the fluorescence energy measured was higher than the abdomen tubes. Furthermore, it is noticeable in Figure 2a that when taking into account the standard deviation, there is a complete overlap between the subcutaneous and the abdomen errors. In addition, the results are based only on measurements of 4 tubes from each type. It may also be noticed in Figure 1 that both types of tubes showed overestimation in some cases and underestimation in others. Therefore, we assume that the negative mean error in the subcutaneous tubes and the positive mean error in the abdomen tubes do not necessarily imply a consistent underestimation or overestimation, correspondingly.
On the other hand, when looking at the mean absolute error, as presented in Figure 2b, we do see a trend of higher absolute errors in the abdomen tubes compared to the subcutaneous tubes. The mean absolute error for the abdomen tubes was 23.13% ± 17.14% and for the subcutaneous tubes: 16.59% ± 11.73%. This is also noticeable in Figure 1, where the star symbols representing the subcutaneous data are all in the range of about 25% error, while two out of four circle symbols representing the abdomen data have errors of about 35%. The higher absolute errors in the fluorescence measurements taken from the abdomen tubes are not surprising and are actually expected. As mentioned earlier, the tissue overlying the fluorescent tubes is thicker for the tubes implanted in the abdomen. The thick tissue causes more absorption and scattering events thus increasing the effective attenuation. As a result, the intensity of the fluorescence signal decreases and the signal to noise ratio decreases too, resulting in higher errors. In addition, in some cases, the fluorescence intensity was so low, that it was very difficult to locate the tube and measure the fluorescence signal. Figure 2. (a) The mean error of estimated concentration ratio (circle symbols) with their 90% confidence interval, for the abdomen tubes (purple) and subcutaneous (blue); (b) the mean of absolute errors of estimated concentration ratios (circle symbols) with their 90% confidence interval, for the abdomen tubes (purple) and subcutaneous (blue).

Sensitivity and Specificity
Several studies have defined a Aβ42:P-tau ratio cutoff value to differentiate between patients developing AD and patients with stable MCI or with other dementias. For example, Buchhave et al. [4] used a cutoff value of Aβ42:P-tau < 6.16 and predicted the future development of AD in MCI patients within 9.2 years (range 4-12 years) with a sensitivity of 88% and a specificity of 90%. The same group used in an earlier study [3] a cutoff value of Aβ42:P-tau < 6.5 for predicting the future development of AD in the same patients within 5.2 years (range 4-6.8 years). These cutoff values may be used as an example for calculating the specificity and sensitivity of our method. The implanted fluorescent tubes were divided according to their true Cy7.5:Cy7 concentration ratio, to two groups: over/under the ratio cutoff value. These two groups were created once with the lower cutoff value (<6.16), which is a cutoff value for early AD diagnosis (9.2 years) and with the second cutoff value (<6.5), which is the cutoff value for a less early AD diagnosis (5.2 years). An estimated concentration ratio was received for each fluorescence ratio measurement of each tube and compared to the cutoff value. A total of 41 estimated concentration ratios were compared to the cutoff value and used for dividing 41 data points to two groups-over/under the cutoff value. These 41 values were compared to their corresponding true values of the concentration ratio. The data points were divided to 4 groups: TN (true negative), FN (false negative), TP (true positive) and FP (false positive). The results according to each cutoff value are presented in Table 2. For both cutoff values the sensitivity was 100%, meaning that all simulated AD 'patients' were diagnosed successfully. The specificity for the higher cutoff value was 82% and for the lower cutoff value was 88%. This result is due to an underestimation in one of the abdomen tubes, which resulted in some FP cases. Obviously, there were more FP cases for the higher cutoff value than for the lower value, resulting in a better specificity in the lower cutoff value.

Time-Resolved Results
Our proposed method involves with fewer risks and side effects than the lumbar puncture process required for CSF collection and for the measurement of amyloid-beta and tau concentration ratio. However, the method still requires an insertion of a needle to the epidural fat in the lumbar section. A feedback system for the needle distance from the CSF can assist in reducing almost completely any complications and risks involved in such a procedure.
Fluorescence lifetime measurements, which may be taken simultaneously with the fluorescence intensity measurements, may provide the basis for such a feedback system. The fluorescence intensity time-resolved curve is affected both by the intrinsic fluorescence lifetime value and by the photon propagation time. The propagation time depends on the tissue optical properties and on the physical distance between the fluorophore and the source and detector. Since the shape of the time-resolved curve is affected by the distance between the fluorophore and the source and detector we can utilize this curve in order to extract the distance of the needle from the CSF. The extracted distance will provide physicians with real-time information about the needle location, thus avoiding penetration of the dura matter.
In our previous report [31] we demonstrated in vitro the effect of variations in the thickness and scattering coefficient of the overlying phantom layer on two features extracted from the fluorescence intensity time-resolved curve. The first feature is t max , which represents the time in which the intensity reached its maximal value and is related directly to the photon path-length. The value of t max increases with an increase in the thickness of the overlying medium due to the longer photon path-length required for the fluorescent photons reach the detector. The second feature is the absolute value of the slope of the decaying part of the intensity curve, in logarithmic scale, which decreases with an increase in the thickness of the overlying medium. In other words, there is a steeper decline in the intensity in thin phantoms or tissues compared to the thicker phantoms or tissues. When the thickness of the overlying media increases, there are more scattering events causing more photons arriving at later times, thus reducing the absolute slope of the intensity decay curve.
The time-resolved curves were measured using the TCSPC system during the animal experiments. The fluorescent tubes implanted in the abdomen were located deeper than the tubes implanted subcutaneously. Since the photon path length of the abdomen tubes is bigger than in the subcutaneous tubes, t max is expected to be higher too. Figure 3 presents the time-resolved fluorescence curve measured from the two tubes implanted in one of the mice: the abdomen tube and the subcutaneous tube. The time-resolved curves were normalized in order to emphasize the difference in t max . As expected, the time in which the abdomen fluorescence reaches its maximal value is higher than the time in which the subcutaneous fluorescence reaches the maximum value: t max (abdomen) = 1.28 ns and t max (subcutaneous) = 1.15 ns.
The absolute value of the slope of the intensity decay curve, in logarithmic scale was also calculated for both cases. For the subcutaneous tubes the absolute value of the slope was 0.75 [1/ns], and for the abdomen tubes the absolute slope was 0.74 [1/ns]. As expected, the absolute value of the slope decreases with an increase in the thickness of the tissue overlying the fluorescent tubes.

Discussion
A method for early detection of Alzheimer's disease using fluorescent probes in the CSF was previously presented by us [31] and validated in vitro and ex vivo. We have further tested and verified the method's accuracy and performance in animals, in this paper. Although the fluorescent probes were not injected directly to the blood and measured in the CSF as the method proposes, the experiments in this study were an additional important step towards testing the method in an AD animal model. Here we implanted tubes with fluorescent samples into two locations-subcutaneous and in the abdomen. The tubes with the clear fluorescent sample simulate the clear CSF layer, and the overlying tissue simulates the epidural fat layer. Since the abdomen tissue of the mouse consists mainly of fatty tissue, we can assume that the optical properties of the abdomen are similar to the epidural fat in the lumbar section. The experiments were performed also on a live tissue with blood flow and other vital signs, therefore making this a good model for the next in vivo step.
The accuracy of the estimation of the ratio between two fluorescent probes was tested for a large range of ratios using these experiments, with an average estimation error of 19.9%. In our previous experiments [31], performed in vitro on polyester phantoms and ex vivo on chicken breast tissue, the average estimation error was lower. It was less than 5% for the in vitro and 11% for the ex vivo experiments. One can notice that there is an increase in the mean error with an increase in the complexity of the overlying medium. This result indicates that the optical properties of the reference phantom, which is used for creating the calibration curve, are not sufficiently close to the properties of a live tissue. The reference phantom is made of polyester resin with titanium dioxide, thus its scattering [33] is similar to the scattering of human tissue, which has a rather flat scattering spectra in the NIR range [34]. On the other hand, the absorption coefficient of human tissue varies in the NIR range with some peaks in the 700-800 nm range. These absorption peaks do not exist in the polyester absorption spectra, and therefore the polyester reference phantom differs in its absorption spectra from a living tissue, in mice or human. We can assume that chicken tissue has more features in its absorption spectra that are similar to human spectra, and obviously living mice are even more similar. The concentration ratio estimation relies on the fact that the optical properties variation in the 700-800 nm range is negligible, and the energy attenuation due to photon propagation is identical in these wavelengths. Therefore, there may be inaccuracies when calculating the ratio based on a calibration curve, which was created using a medium with a different absorption spectrum than the measured tissue. Nevertheless, the errors are still rather small and reasonable. The system's accuracy may be improved using a different reference phantom, which better resembles the human tissue optical properties, for instance by adding an absorber, such as blood to the polyester. Such an absorber may assist in creating an absorption spectrum similar to the human tissue spectrum.
The fluorescent probes that were implanted subcutaneously had slightly lower estimation errors compared to those implanted deeper in the abdomen. This was explained by the low fluorescence intensity in the abdomen tubes, caused by a thicker overlying tissue. The decrease in the fluorescence intensity caused a lower signal to noise ratio and bigger errors. In addition, the concentration ratio estimation is based on the assumption that the measured fluorescence intensity ratio is affected mainly by the concentrations of the fluorophores and that the energy attenuation due to photon propagation is identical in both wavelengths. However, there are some differences in the absorption and scattering coefficients between the wavelengths used for the excitation of the two fluorophores, and therefore the errors in the concentration ratio estimation increase with an increase in the tissue thickness. Since the abdomen tubes have thicker overlying tissue than the subcutaneous tubes, the inaccuracy that results from not taking into account photon propagation attenuation has a higher effect, causing bigger errors. There are several actions that may be helpful in avoiding high errors in thick tissue. The first is obviously to avoid performing measurements when there are high distances between the needle and the CSF. This may be done by slow insertion of the needle until a sufficiently high fluorescence signal is measured, with an acceptable signal to noise ratio. In addition, we presented the option to use the fluorescence time-resolved measurements for estimating the distance between the needle and the CSF, and the scattering coefficient. These two estimated values may be used as a correction factor that takes into account the tissue thickness and its optical properties without totally neglecting the photon propagation attenuation.
Although in some cases, the ratios were estimated with an error of 35%, the sensitivity and specificity still remained high. This may be explained by the fact that for low concentration ratios, an error of 35% still resulted in a relatively low difference between the true and estimated ratio. For instance, in one of the cases, the true ratio was 3.57 and the estimated ratio was 4.86, which is equivalent to 36% error. However, the difference between the true and estimated ratio is 1.29, a rather low difference that will not cause false predictions. In the high concentration ratios, the big errors caused a larger difference between the true and estimated ratio values, but the cutoff value was relatively distant from the true ratio. As a result, in those cases, even big errors did not cause too many false predictions. On the other hand, the concentration ratio values that were close to the cutoff value were estimated with a very high accuracy so the sensitivity and specificity remained high. We can assume that our method requires high accuracy in ratios that are close to the cutoff ratio but can afford bigger errors in the extreme cases of low and high concentration ratios since big errors in these values will not result in false predictions.
The fluorescence properties of Cy7 and Cy7.5 probes are not identical to one specific probe developed for AD (presented in Section 1). Cy7 has absorption/emission peaks at 750/773 nm, an extinction coefficient of 199,000 [M −1 cm −1 ] and quantum yield of 0.3. Cy7.5 has absorption/emission peaks at 788/808 nm and an extinction coefficient of 223,000 [M −1 cm −1 ]. When comparing the fluorescent properties of the AD probes we see some differences in the absorption/emission peaks but similar orders of magnitudes of the extinction coefficient and quantum yields. For example, AOI987, which was used for imaging Aβ, has an absorption/emission peaks at 650/670 nm, an extinction coefficient of 64,570 [M −1 cm −1 ], and a quantum yield of 0.4. THK-265 showed similar properties-absorption/emission peak at 630/650 nm, extinction coefficient of 96,198 [M −1 cm −1 ], and quantum yield of 0.38. CRANAD-58 has a similar absorption peak in 630 nm but an emission peak in 750 nm in the presence of Aβ. For tau imaging, CyDPA2 was presented with an absorption/emission peaks at 826/833 nm, extinction coefficient of 113,553 [M −1 cm −1 ] and quantum yield of 0.049. The differences in the absorption/emission spectra between the fluorescent probes used in this study and the AD probes may affect the accuracy of the concentration ratio estimation. If the changes between the absorption spectra of the Ab and tau probes are big, the assumption that the optical properties are similar is less accurate. On the other hand, if the two probes have completely separate spectra, the possibility of unwanted excitation of the tau probe when one wants to excite only the Ab probe will be significantly reduced. This phenomenon is one of the problems we had to deal with during our experiments. The absorption spectra of Cy7 and Cy7.5 partly overlapped, causing excitation of both probes together when we needed to measure the fluorescence of each one of them separately. Despite the use of filters for fluorescence emission collection, the combined excitation increased the intensity of the measured fluorescence and possibly contributed to some of the errors.
One of the main advantages of the proposed method, compared to other methods, which measure concentrations in vivo, is that it does not require a priori knowledge regarding the tissue anatomical and optical properties. It is true that measuring the concentrations of amyloid beta and αtau directly from a collected CSF sample will probably result in a better accuracy. However, this will require performing a lumbar puncture procedure. In our method, as opposed to the lumbar puncture, there is no need to penetrate the dura and collect CSF samples, and the risks and side effects are significantly reduced.
These risks may be further reduced by adding a feedback system to our method, which will provide real-time information regarding the location of the needle and its distance from the dura matter and the CSF. This feedback system, as we have shown, is based on the time-resolved intensity curves, by extracting the values of t max and the slope, thus enabling accurate estimation of the absolute value of the needle distance from the CSF. The accuracy of the estimation may be increased by taking multiple time-resolved intensity measurements during the insertion of the needle into the lumbar section. The needle distance may then be estimated without a priori knowledge of the tissue's optical properties.
In addition, once a fluorescence signal is detected, with a sufficient intensity that ensures a good signal to noise ratio, further insertion of the needle may be stopped. The system will evaluate the concentration ratio and the needle distance from the CSF and may indicate to the physician how to continue. If the distance is big enough and the signal to noise ratio is still poor, the system will indicate to the user to continue to push in the needle. In case that the estimated distance between the needle and the CSF is too small the system will indicate the physician to stop additional insertion in order to reduce the risk of dura penetration. We demonstrated the ability to accurately estimate the concentration ration of two fluorescent probes underlying a tissue with a thickness of up to 5 mm. Therefore, we assume that the system may alert the user of being close to the dura in an estimated distance of 3 mm, thus avoiding penetration of the dura even with distance estimation errors of several hundreds of percent but with estimating the concertation ratio with a rather high accuracy.

Conclusions
We have demonstrated the possibility to accurately estimate the concentration ratio of two fluorescent probes implanted several millimeters under the skin in living mice. This study has a higher complexity than our previous study, which tested the estimating accuracy in vitro and ex vivo. Here the concentration ratio estimation takes into account the absorption and scattering spectra of an in vivo tissue, pulsatile blood flow and respiration movements. It is a preparation for in vivo studies that will be performed on AD mice models with markers injected to the blood and reaching the CSF. The proposed method may estimate in the future the CSF Aβ42:P-tau ratio without the need to penetrate the dura matter and collect CSF samples. It does not require any a priori information regarding the anatomical and optical properties of the lumbar tissue. Considering these advantages, this technique could be used as an annual test, which may assist in early AD diagnosis and increase the efficacy of treatments. In recent years there are reports on additional potential CSF bio-markers that may be used as early detection [35,36]. Our method is generic and only relies on having a fluorescence marker that can be attached to those disease markers. Therefore, our proposed method can also function on other Alzheimer's disease markers, so it will be even more powerful than relying on the ratio of the two markers that we have discussed in this work.