Improved in vivo delivery of m-THPC via pegylated liposomes for use in photodynamic therapy

Journal of Controlled Release 157 (2012) 196–205

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Improved in vivo delivery of m-THPC via pegylated liposomes for use in photodynamic therapy Melissa J. Bovis a,⁎, Josephine H. Woodhams a, Marilena Loizidou b, Dietrich Scheglmann c, Stephen G. Bown a, Alexander J. MacRobert b a b c

National Medical Laser Centre, Division of Surgery and Interventional Science, University College London, London, UK Department of Surgery, Division of Surgery and Interventional Science, University College London, London, UK Biolitec AG, Jena, Germany

a r t i c l e

i n f o

Article history: Received 27 May 2011 Accepted 22 September 2011 Available online 1 October 2011 Keywords: m-THPC Photosensitiser Photodynamic therapy Liposomes Polyethylene glycol Pharmacokinetics

a b s t r a c t Pegylated liposomal nanocarriers have been developed with the aim of achieving improved uptake of the clinical PDT photosensitiser, m-THPC, into target tissues through increased circulation time and bioavailability. This study investigates the biodistribution and PDT efficacy of m-THPC in its standard formulation (Foscan®) compared to m-THPC incorporated in liposomes with different degrees of pegylation (FosPEG 2% and FosPEG 8%), following i.v. administration to normal and tumour bearing rats. The plasma pharmacokinetics were described using a three compartmental analysis and gave elimination half lives of 90 h, 99 h and 138 h for Foscan®, FosPEG 2% and 8% respectively. The accumulation of m-THPC in tumour and normal tissues, including skin, showed that maximal tumour to skin ratios were observed at ≤24 h with FosPEG 2% and 8%, whilst skin photosensitivity studies showed Foscan® induces more damage compared to the liposomes at drug-light intervals of 96 and 168 h. PDT treatment at 24 h post-administration (0.05 mg kg−1) showed higher tumour necrosis using pegylated liposomal formulations in comparison to Foscan®, which is attributed to the higher tumour uptake and blood plasma concentrations. Clinically, this improved selectivity has the potential to reduce not only normal tissue damage, but the drug dose required and cutaneous photosensitivity. © 2011 Elsevier B.V. All rights reserved.

1. Introduction Photodynamic therapy (PDT) is a well established anticancer treatment that uses a light-activated photosensitising agent to induce necrosis mediated by the production of reactive oxygen species [1]. It is a minimally invasive technique that is repeatable and unlike conventional cancer therapies elicits no major systemic toxicity. Moreover, this process only occurs upon illumination of the non-toxic photosensitiser (PS) at an appropriate wavelength which renders it selective to the light treated area. Despite these favourable properties, there are problems associated with the biodistribution of PS to normal tissues and its clearance, resulting in residual cutaneous photosensitivity which must be managed until the drug is eliminated [2]. The selectivity of tumour uptake with PDT photosensitisers is also marginal and improvements are currently being sought.

Abbreviations: PDT, photodynamic therapy; PS, photosensitiser; m-THPC, 5,10,15,20tetra-(m-hydroxyphenyl)chlorin; PEG, polyethylene glycol; RES, reticuloendothelial system; EPR, enhanced permeability and retention; DPPC, dipalmitoylphosphatidylcholine; DPPG, dipalmitoylphosphatidylglycerol; MC28, methylcholanthrene-induced fibrosarcoma cells; HL, Hooded Lister; DLI, drug-light interval. ⁎ Corresponding author. Tel.: + 44 20 7679 9069; fax: + 44 20 7813 2828. E-mail address: [email protected] (M.J. Bovis). 0168-3659/$ – see front matter © 2011 Elsevier B.V. All rights reserved. doi:10.1016/j.jconrel.2011.09.085

The chlorin investigated here, m-THPC, is a highly potent, second generation PS that exhibits several favourable characteristics for PDT[3] and has been approved for the treatment of head and neck cancers in the EU and Japan. Being a chlorin it exhibits stronger absorption than porphyrins at longer wavelengths (652 nm); however, m-THPC is hydrophobic and is prone to aggregation, which presents problems in optimising its formulation [4]. The aggregated form of m-THPC is less photoactive and binds strongly to serum proteins [5]. Rapid uptake of aggregates by cells of the reticuloendothelial system (RES) may lead to higher inter-patient variability of m-THPC accumulation in tumour tissue with the current micellar formulation, Foscan® [6]. The unique microenvironment of solid tumours also limits PS delivery and distribution within cancerous tissue [7], further contributing to differences in PDT response [8]. In PDT, liposomes have been used in recent years due to their suitability for packaging large quantities of hydrophobic photosensitisers, including m-THPC into their lipid shells [9,10] and their ability to accumulate in tumour tissue through the enhanced permeability and retention (EPR) effect as a result of their size [11]. However, liposomes are still susceptible to recognition by the host's immune system and rapid uptake into the RES [12]; therefore polyethylene glycol (PEG), a biocompatible polymer, has been used to coat the outer surface of the liposome. This coating sterically stabilises the liposome and increases its hydrophilicity to minimise the recognition and binding of opsonins [13], thereby hindering the loss of the liposome from circulation [14].

M.J. Bovis et al. / Journal of Controlled Release 157 (2012) 196–205

Consequently, this formulation should lead to improved bioavailability of m-THPC. Pegylation is further believed to inhibit the release of hydrophobic compounds localised within the liposomal membrane by minimising the adsorption of serum proteins [15]. Varying, but defined, degrees of liposome pegylation were used in this study to assess the effect of increased pegylation on m-THPC biodistribution. In a previous veterinary study using cats, pegylated liposomal m-THPC uptake was studied in the plasma, tumour and skin, but the degree of pegylation (2.5–5%) was not precisely defined [10]. In this study using rats, pegylations of 2% and 8% (molar equivalent) were chosen as the lower and upper limits since higher percentages of PEG are known to disrupt the integrity of the liposome membrane [16]. Recently, in a joint study including our laboratory, the in vitro and photophysical properties of the same liposomes were investigated [17]. Fluorescence lifetime data suggest m-THPC is aggregated when incorporated in the lipid shell of the liposomes since fluorescence is strongly quenched, however this is not necessarily a serious drawback, as following cellular uptake, the liposome is rapidly degraded and m-THPC is then released into its monomeric photoactive form inside the cell. The in vitro results showed that A549 lung carcinoma cells were killed very efficiently under illumination, following incubation with the pegylated liposomal m-THPC formulations. The aim of the present study was to investigate the in vivo biodistribution and accumulation of two pegylated liposomal m-THPC formulations (FosPEG 2% and FosPEG 8%), in comparison to standard Foscan® (m-THPC). This was achieved through pharmacokinetic analysis in both normal Wistar rat and subcutaneous syngeneic fibrosarcoma (MC28) Hooded Lister rat models, which have been used successfully in previous m-THPC (Foscan®) studies [18]. The uptake of m-THPC through these pegylated liposomes was then correlated with measurements of PDT efficacy and skin photosensitivity.

197

Female Wistar rats (180–220 g) were used for normal tissue pharmacokinetic studies and skin photosensitivity studies, since skin pigmentation is minimal. The same model was used previously by Wilson et al. in studies with a different photosensitiser [19]. Each formulation was administered intravenously via a tail vein injection at a dose of 0.3 mg kg−1 m-THPC. At a specified time point between 2 and 168 h, animals were killed by cervical dislocation. For plasma readings, animals were killed immediately after injection at an estimated time of ≤5 min. A methylcholanthrene-induced fibrosarcoma cell line (MC28), syngeneic and transplantable to Hooded Lister (HL) rats was cultured in accordance with previous studies [18,20]. Female HL rats (150–220 g) were inoculated subcutaneously in the lower flank with approximately 1–2 × 10 6 MC28 cells. Tumours were monitored continuously and reached an optimal size of approximately 10 mm 3 after 7–10 days. 2.3. Pharmacokinetic study

Biolitec AG (Jena, Germany) kindly provided all m-THPC formulations. Foscan® was supplied in its standard formulation (m-THPC in ethanol/propylene glycol) at a stock concentration of 4 mg mL −1. Liposomal formulations of m-THPC, FosPEG 2% and FosPEG 8% were prepared as a 9:1 mixture of dipalmitoylphosphatidylcholine (DPPC) and dipalmitoylphosphatidylglycerol (DPPG). The degree of pegylation was 2% and 8% (molar equivalent ratio) using 1,2-distearoyl-sn-glycero-3phosphoethanolamine-N-[amino (polyethylene glycol)-2000] (DSPEPEG2000). (Further liposome preparation details can be found in Ref. [17].) These were provided as stock solutions containing a molar equivalent concentration of 2.21 mM (or 1.5 mg mL−1) m-THPC in 10 mM histidine buffer, containing 50 mg mL−1 glucose at pH 6.5. The mean number of m-THPC molecules was estimated at 2 × 10 4 per liposome, based on the 9:1 molar ratio of lipid to m-THPC. Mean particle size distribution and particle characterisation were assessed using photon correlation spectroscopy, differential scanning calorimetry and cryo-TEM. The mean particle diameter of liposomal formulations was 120 nm. All formulations remained stable in size for up to 12 months in storage buffer and upon dilution, with no drug precipitates or aggregates observed.

Tissue samples selected for pharmacokinetic analysis included muscle, skin (right abdominal wall), liver, spleen, kidneys, lung, blood plasma and tumour, taken from three animals at each time point. Immediately post-mortem tissues were removed under subdued lighting at 2, 4, 6, 18, 24, 72, 96 and 168 h from normal Wistar rats after intravenous administration of each m-THPC formulation at 0.3 mg kg −1 and 2, 4, 6, 24 and 72 h for HL tumour rats, due to maximal tumour growth threshold being attained. Blood samples (~ 3 mL) were also taken at an additional ≤ 5 min interval and centrifuged to separate the plasma at 2000 rpm for 10 min for pharmacokinetic analysis. Negative control animals at 0 h were run simultaneously. Samples were stored in the dark at − 80 °C postmortem. m-THPC accumulation in tissues was measured through a chemical extraction method combined with spectrofluorimetric analysis of m-THPC in the extract, adapted from Kascakova et al. [21]. Tissue samples of approximately 0.1 g wet weight were run in triplicate. These were incubated in 2 mL of Solvable™, an aqueous-based alkaline solvent (Perkin-Elmer, UK), for approximately 2 h at 50 °C in a shaking water bath, until completely dissolved without any visible tissue residue. Fluorescence detection was used to construct a linear standard curve of known m-THPC concentrations (0–5 μM) from control tissues for each organ, prepared under identical conditions and diluted in the same solvent (Solvable™) as test samples. The solubilised tissues were aliquoted (300 μL) into a 96 well plate. The fluorescence signal from each well was detected using a Perkin-Elmer LS 50B fluorescence spectrometer (Perkin-Elmer, UK) linked to a plate reader. Front surface excitation/detection was employed using a fibre-optic probe which minimises reabsorption and polarisation artefacts. Excitation and emission wavelengths for m-THPC fluorescence measurements were set at 423 nm and 652 nm respectively. Fluorescence readings were used to calculate the mean tissue concentration of m-THPC. Readings from negative control samples (without m-THPC) were deducted to correct for the autofluorescence of each organ. The mean and the standard deviation (SD) for all three animals at each time point were calculated. The plasma pharmacokinetics were analysed by compartmental and non-compartmental mathematical methods [22,23].

2.2. Animals and tumour models

2.4. Skin photosensitivity studies

All animal experiments were carried out under the authority of project and personal licences granted by the UK Home Office and with reference to NCRI (National Cancer Research Institute) guidelines for the Welfare of Animals in Experimental Neoplasia (2010). All procedures were performed under general anaesthesia with inhaled isofluorane. Buprenorphine hydrochloride was given subcutaneously for postoperative analgesia where necessary.

Skin photosensitivity studies required shaving, depilating (Neet™) and cleaning an area on the lower flank of female Wistar rats (180– 220 g). Following m-THPC administration (i.v. 0.3 mg kg−1), 4 circular areas of skin measuring 0.5 cm in diameter and 1 cm apart (grid design) were exposed to either 0 (0 J), 5 (30 J), 15 (90 J) or 30 min (180 J) of 100 mW cm−2 of light (1 sun = solar equivalent spectrum) sequentially from a solar simulator source (Olympus CLV-S30, Cambridge, UK) at

2. Materials and methods 2.1. Photosensitisers

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drug-light intervals (DLI) of 96 and 168 h. This system incorporated a 300 W (Olympus Endoscopic) xenon lamp, providing a uniform 2 cm diameter beam at the skin surface. Surrounding skin tissue was protected from light using a light impenetrable fabric. Illuminated areas were marked adjacently. Light power was measured using a laser power metre (Gentec, UK) and was calibrated before and after each treatment. Treatments were carried out in a climate controlled room to prevent overheating of the skin. Visual assessment of skin reactions were carried out 24 h post-treatment and digital images taken. Imaged areas were then assessed by two independent scorers using a blinded grading model, previously established by Weersink et al. [19]. Grading was carried out on all images at one time to ensure subjective analysis across the data set was relative. Each m-THPC formulation was allocated an ordinal value (0–8) and scores were averaged over each exposure time. Areas of skin treated at 0 min (no light control) and 30 min were removed for histological analysis. A separate group of animals that received no PS was also treated with 30 min of light to ensure skin effects were a result of the presence of PS, not light exposure.

Concentration of m-THPC (µg ml-1)

198

10 Foscan FosPEG 2% FosPEG 8%

1

0.1

0.01

0

24

48

72

96

Time (h) Fig. 1. Semi-log plot of mean m-THPC concentration in blood plasma after intravenous administration of 0.3 mg kg−1 m-THPC in 'standard' Foscan or FosPEG 2% and FosPEG 8% liposomal formulations into female Wistar rats. This plot is fitted according to a three compartmental model. Data points show the mean± SD, n = 3.

2.5. Photodynamic therapy on MC28 tumours

2.7. Statistical analysis

Treatment with each m-THPC formulation was initiated when tumours had reached an optimal diameter of 10 mm. A DLI of 24 h was chosen for PDT studies based on chemical extraction data of tumour tissue. Clinical m-THPC doses of 0.3 mg kg −1 (high) and 0.05 mg kg −1 (low) [24] were administered to two groups of animals. Tumours were irradiated with red laser light interstitially from a 652 nm diode laser (Diomed, Cambridge, UK) using a 400 μm barecleaved tip optical fibre inserted approximately 1 mm into the tumour capsule via a small incision in the overlying skin and in the MC28 tumour capsule. This irradiation method has been used in our previous studies with m-THPC in the same tumour model [18] and mimics interstitial clinical PDT with Foscan, predominantly carried out on head and neck tumours whereby the laser fibre is inserted into the tumour. A total of either 2 J or 10 J of light at 100 mW cm −2 (100 s) was delivered to each tumour. Each treatment group consisted of five animals. Animals were killed 24 h after treatment by cervical dislocation and whole tumours resected for histological analysis.

Mean and standard deviation was calculated for each animal group (±SD, n = 3–5). All data were represented as mean ± SD. Statistical analysis was carried out using a two-tailed Student's t-test and a Mann–Whitney t-test for PDT data. p ≤ 0.05 was considered statistically significant, unless stated otherwise. 3. Results 3.1. Plasma pharmacokinetics of m-THPC (Foscan® versus FosPEG 2% and 8%) Fig. 1 shows the plasma clearance after a single i.v. injection of 0.3 mg kg−1 m-THPC in Foscan, FosPEG2% and 8% formulations into the normal Wistar rat model between ≤5 min and 96 h. m-THPC plasma concentrations peaked at the earliest time point for each formulation and appeared to decline exponentially. However, m-THPC clearance from plasma fitted multiple exponential decays; therefore data were analysed using both a compartmental and non-compartmental approach. Using a compartmental model, the data best fit a tri-exponential decay, which gave

2.6. Histology and measuring necrosis (Hamamatsu Nanozoomer) Whole tumours and skin tissue samples removed post-mortem were immersed in 4% neutral formalin buffer (4% w/v formaldehyde in phosphate buffered saline) for a minimum of 24 h at 20 °C. Samples were processed by routine methods. The tumours were cut in half (parallel to the laser fibre) and skin samples were cut through the centre of the treatment area, and adjacent halves of the tissue were embedded face down in paraffin wax blocks. Four-micrometre sections were cut and mounted on Vectabond (Vecta laboratories, UK) treated glass slides. Three sections were taken from each of the tissue halves. Slides were stained with Harris haematoxylin and eosin. Whole slides were scanned with the Hamamatsu Nanozoomer (Hamamatsu Photonics UK Ltd). Hamamatsu virtual microscopy imaging software was used to observe markers indicative of skin photosensitivity i.e. erythema and oedema. PDT damage in tumour tissue was assessed by measuring the surface area of necrosis from each section, as used previously in m-THPC PDT [25]. The damage was calculated from each tumour as a percentage of the whole tumour surface area. Six sections were averaged per tumour (3 from each tumour half). The mean percentage surface area of tumour necrosis was calculated through blind analysis per group of five identically treated animals and all data were represented as mean± SD.

Table 1 Plasma pharmacokinetic parameters of m-THPC after i.v. injected dose of 0.3 mg kg−1 m-THPC in Foscan and each liposomal formulation in Wistar rats, calculated using the exponential equations of the three compartment model (Ae_αt + Be_βt + Ce_γt), and the non-compartmental method. Pharmacokinetic parameter

Foscan

FosPEG 2%

FosPEG 8%

Three compartmental model Initial dosage (Do, mg kg−1) Initial concentration (Co, μg mL−1) Initial volume of distribution (Vd, mL kg−1) Vd of first compartment (mL kg−1) Vd of second compartment (mL kg−1) Vd of third compartment (mL kg−1) t1/2 of first compartment (h−1) t1/2 of second compartment (h−1) t1/2 of third compartment (h−1)

0.3 0.7 407.4 1373.9 640.1 6086.2 0.9 3.3 90.0

0.3 8.7 34.6 71.1 69.5 2231.8 1.0 2.3 99.0

0.3 16.9 17.7 21.8 100.9 1505.8 0.6 5.5 138.6

Non-compartmental model Plasma clearance (Cl, mg kg−1 h−1) Mean residence time (MRT, h−1) Volume of distribution (Vd, mL kg−1) Half life (t1/2, h−1) Elimination rate constant (Kel, h−1)

30.0 61.9 1875.7 42.9 0.016

6.41 72.5 464.2 50.2 0.014

3.21 90.3 289.5 62.6 0.011

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three half-lives and elimination rate constants for Foscan, FosPEG 2% and 8% (Table 1). The biological half-life (t1/2) of the third compartments (log/linear phase) show a 1.5-fold increase between Foscan and FosPEG 8%, 90 h vs. 138.6 h, which is confirmed by figures obtained through non-compartmental analysis, 42.9 h vs. 62.6 h (Table 1). The volume of distribution (Vd) for FosPEG 2% and 8% is much lower than Foscan (Table 1), suggesting all formulations accumulate in different tissues but FosPEGs are more confined to the vasculature. 3.2. Bio-distribution of m-THPC (Foscan®, FosPEG 2% and 8%) In normal tissues of the Wistar rat, the highest m-THPC concentrations were observed in the highly perfused tissues, such as the liver, spleen and

Concentration of m-THPC (µg g-1)

4

Foscan

4

FosPEG 2% FosPEG 8%

3

3

2

2

1

1

0

0 0

20

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160

C Concentration of m-THPC (µg g-1)

lungs (Fig. 2A–C). Concentrations of m-THPC peak at the earliest time points, between 2 and 4 h for all formulations, ranging from approximately 2.5±3.5 μg g−1 in the liver (Fig. 2A), and 1.0±2.5 μg g−1 in the spleen (Fig. 2B). FosPEG 2% shows complete clearance from both organs by ~96 h. FosPEG 8% displays slower distribution kinetics and in lung tissue remains consistently higher than Foscan (~6 fold) and FosPEG 2% (~2– 3 fold) over 168 h, which are both cleared by this time (Fig. 2C). mTHPC levels in the kidney (Fig. 2D) plateau from 2 to 168 h for all formulations, reaching concentrations of approximately 0.7 μg g−1 for Foscan and FosPEG 2% but not exceeding 0.2 μg g−1 for FosPEG 8%. No clear differences between m-THPC formulations are observed in muscle (Fig. 2E), whereas skin data show significantly elevated concentrations of FosPEG 8% from 96 h compared to Foscan and FosPEG2%, which have similar kinetic profiles patterns (Fig. 2F). Both the muscle and skin accumulate

B

A

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D 3.0

3.0

2.5

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1.5

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1.0

0.5

0.5

0.0

0.0 0

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E Concentration of m-THPC (µg g-1)

199

F 0.5

0.5

0.4

0.4

0.3

0.3

0.2

0.2

0.1

0.1

0.0

0.0

0

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Time (h)

120

140

160

Time (h)

Fig. 2. Concentration of m-THPC in selected tissues of the Wistar rat as a function of time following an intravenous injection of 0.3 mg kg−1 m-THPC in standard Foscan, FosPEG 2% and FosPEG 8% formulations. (A) liver, (B) spleen, (C) lung, (D) kidney, (E) muscle and (F) skin. Data points show the mean ± SD, 3 ≤ n ≤ 4.

200

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the least amount of m-THPC for all formulations and an elimination phase did not occur for these peripheral tissues in the time course of this study.

the epidermis and dermis respectively. No photosensitive effects to skin are detected at 0 min of light exposure (Fig. 4) and no damage was observed with those animals that were exposed to light but did not receive m-THPC (data not shown).

3.3. Skin photosensitivity studies with m-THPC (Foscan®, FosPEG 2% and 8%) 3.4. Tumour accumulation of m-THPC (Foscan®, FosPEG 2% and 8%) A positive correlation was identified between the length of light exposure (0 to 30 min) and observable photosensitive effects from graded skin analysis (Fig. 3). Visually, Foscan evoked greater cutaneous photosensitivity over time in comparison to milder effects observed with FosPEG 2%, which displays the lowest levels of skin photosensitivity at both DLIs at 30 min exposure. FosPEG 8% demonstrates a slightly higher degree of skin photosensitivity at 96 h in comparison to both Foscan and FosPEG 2%, as a result of greater skin damage observed between 15 and 30 min light exposure. This supports chemical extraction data of FosPEG 8% in skin tissue at this time (Fig. 2F). However by 168 h, Foscan scored highest at a light exposure time of 30 min. No blistering or necrosis of skin tissue was observed visually (Score 8) with any m-THPC formulations at the dose used in this study. Control tissue exposed to 0 min light was unaffected (Score 0), as were animals that only received light. Fig. 4 illustrates typical histological rat skin photosensitivity from the same treatment groups at 0 min (control) and 30 min light exposure. Qualitative differences exist between m-THPC formulations, with Foscan eliciting extensive superficial damage to the epidermis at DLI of 96 h, illustrated by mass cell death, in comparison to FosPEG formulations which are slightly milder. At a DLI of 168 h, photoinduced damage is reduce with FosPEG 2%, showing no signs of sensitivity, whereas Foscan and FosPEG 8% demonstrate localised damage to regions of 7 96 h 6

168 h

5

3.5. Tumour PDT-response with m-THPC (Foscan®, FosPEG 2% & 8%) All formulations of m-THPC cause significant tumour necrosis, expressed as a percentage of the total tumour tissue (Fig. 6), in comparison to the control treatment group which received no drug (p ≥ 0.001). For each of the three variable treatment groups there is also a significant difference in percentage tumour necrosis between Foscan versus FosPEG formulations (p ≤ 0.001). When exposed to 10 J of light following administration of 0.3 mg kg −1 of m-THPC, percentage tumour necrosis with FosPEG formulations is slightly lower than that with Foscan (Fig. 6). In contrast, with either low dose m-THPC (0.05 mg kg−1) or low light energy (2 J) both FosPEG 2% and 8% produced a significantly greater percentage of tumour necrosis in response to PDT compared to Foscan, by either 10% or 20% (p ≥ 0.001). 4. Discussion

4

Score

There is a significant increase in maximal m-THPC uptake in tumour tissue of the HL rat between 6 and 24 h with FosPEG 2% and 8%, in comparison to Foscan (p ≤ 0.001) of approximately three-fold (~1.14 ± 0.98 vs. 0.33 μg g−1) (Fig. 5A). Although there is no significant difference in accumulation between the two FosPEG formulations over the majority of the time series, FosPEG 8% administration results in more selective retention in the tumour tissue for up to 72 h post injection. At this time period it remains at more than four times the concentration obtained using Foscan (0.62 vs. 0.14 μg g−1). By 24 h, all m-THPC formulations enter a terminal linear elimination phase from tumour tissue (Fig. 5A). Mean concentration ratios of tumour to skin tissue from the HL rat model were calculated for all m-THPC formulations (Fig. 5B) and show the greatest difference in m-THPC uptake between tumour and skin is at ~6 h. FosPEG 2% elicits almost 5 times the concentration ratio found with Foscan, whilst FosPEG 8% is slightly lower.

3

2

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0 0

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15 30

Foscan

0

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15 30

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15 30

FosPEG 2% FosPEG 8%

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30

Fig. 3. Wistar rats exposed to 0, 5, 15 or 30 min of 100 mW cm−2 (1 sun = solar equivalent spectrum) of light from a solar simulator source at drug-light intervals (DLI) of 96 and 168 h, following an intravenous injection of 0.3 mg kg−1 m-THPC in Foscan, FosPEG2% and FosPEG 8%. Visual assessment of skin reactions carried out 24 h post-treatment and imaged areas scored by a double-blind grading model. Scores were averaged over each exposure time for each m-THPC formulation.

The aims of this study were to investigate the biodistribution and accumulation of m-THPC in different tissues of normal and MC28 tumour-bearing rat models when delivered by pegylated liposomal nanocarrier formulations versus standard Foscan. PDT efficacy and skin photosensitivity were additionally compared between m-THPC formulations in relation to their distribution. The plasma pharmacokinetics of Foscan, FosPEG 2% and 8%, can be described by a tri-exponential decay curve (Fig. 1). A bolus intravenous injection of 0.3 mg kg−1 of m-THPC in each formulation produces an instantaneous peak and a rapid distribution into the tissues, representing the first phase of elimination. This is followed by a less rapid m-THPC redistribution and finally, a terminal phase of elimination from tissues [22]. This decay model matches previous studies carried out in murine models with Foscan [26,27] and a nonpegylated liposomal m-THPC formulation, Foslip [9]. The Foscan half-life data obtained for each elimination phase (0.9, 3.3 and 90.0 h) (Table 1) are very close to the data of Jones et al. [26] (0.46, 6.91 and 82.5 h), also in rats. Studies in felines using an earlier FosPEG formulation by Buchholz et al. [10] revealed equal peak plasma concentrations of Foscan (~0.45 μg mL−1) at ≤5 min–2 h, as illustrated here (Fig. 1), but found maximal plasma concentrations of FosPEG were only ~3.5 times higher than Foscan. However, relatively few details were provided on the liposome composition used in that study, and the degree of pegylation (2.5–5%) was not tightly defined; therefore we cannot draw any definitive comparisons with the present study. A smaller volume of distribution was observed through non-compartmental analysis with FosPEG 2% (464 mL kg−1) and FosPEG 8% (290 mL kg−1) in

M.J. Bovis et al. / Journal of Controlled Release 157 (2012) 196–205

96 h

Foscan

201

FosPEG2%

FosPEG8%

M M

D

Ad

HF

0 min (control)

D

D

ED 1.0 mm

500 µm

1.0 mm

D

M

D

D

1. 0 mm

ED

30 min ED ED 100 µm

168 h

FosPEG2%

Foscan 600 µm

1.0 mm

1.0 mm

600 µm

FosPEG8% 1.0 mm

M

M Ad

0 min (control)

D D

D M

HF Ad 1.0 mm

1.0 mm

500 µm

D

D ED

30 min HF ED

D

ED HF

Fig. 4. Histological sections of skin tissue removed from the animal 24 h after light treatment. Skin control samples at 0 min (no light) and treated samples at maximal exposure of 30 min were collected and cut through the centre of the treatment area. Adjacent halves of the skin tissue were sectioned (4 μm) and stained with H&E. Images were observed with the Hamamatsu Nanozoomer. M—muscle, Ad—adipocytes, D—dermis, HF—hair follicle, ED—epidermis.

comparison to studies using Foslip (709 mL kg−1) [9], indicating pegylated liposomal m-THPC nanocarriers are more confined to the vasculature with increased pegylation. By comparison, Foscan has a much larger volume of distribution (1876 mL kg −1) suggesting it preferentially accumulates in some tissues. Likewise, compartmental analysis (Table 1) demonstrates an initial volume of distribution of 407 mL kg −1 with Foscan, which is extremely high given the blood volume of a rat ~ 50–70 mL kg −1. This indicates that the initial retention of Foscan in the plasma compartment is very low as a result of Foscan being rapidly bound by proteins in the blood and taken up by cells of the reticuloendothelial system (RES) [28]. In contrast, FosPEG 2% and 8% have a low initial volume of distribution of 34.6 and 17.7 mL kg −1 respectively. It is believed the PEG polymer coating

sterically stabilises the liposomal particle and prevents rapid binding by scavenger receptors (opsonins) of macrophages at the earliest time points, thus increasing the longevity and accumulation of mTHPC in the blood circulation. The extended m-THPC half-lives of FosPEG 2% (99.0 h) and FosPEG 8% (138.6 h), however, exhibit only moderate differences compared to Foscan (90.0 h) (Table 1). It is possible that by 96 h the liposomes have broken down in vivo so that only free m-THPC is being measured in the terminal phase of elimination, resulting in similar half-lives. However, pegylated liposomes are known to have improved stability in serum in terms of drug release, as demonstrated in a previous study, using a similar pegylated liposomal composition (DPPC:DPPG, 9:1) [29]. A reduction in the release of a hydrophobic drug (paclitaxel) from pegylated

202

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Concentration of m-THPC (µg g-1)

A 1.6

*

1.4

*

1.0

*

0.8 0.6

*

0.4 0.2 0.0 0

Concentration Ratio (Tumour/Skin)

B

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1.2

12

24

36

48

60

72

Time (h) 7 Foscan FosPEG2% FosPEG 8%

6 5 4 3 2 1 0 0

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Time (h) Fig. 5. (A) Concentration of m-THPC in MC28 tumour tissue of female Hooded Lister rat as a function of time following an intravenous injection of 0.3 mg kg−1 m-THPC in standard Foscan, FosPEG 2% and FosPEG 8% formulations. (B) Mean concentration ratio of tumour to skin in female Hooded Lister rats over a time series following intravenous injection of 0.3 mg kg−1 m-THPC in Foscan, FosPEG2%, FosPEG 8% formulations. Data points show the mean ± SD, n = 4.

liposomes was observed over 96 h when incubated at 37 °C in human serum (35% of initial encapsulated drug) compared to non-pegylated liposomes (72%). This reduction could be attributed to the PEG coating, which inhibits serum-induced drug leakage. The same conclusion was

PDT induced tumour necrosis (% area of necrosis)

100 90 80

Foscan FosPEG2% FosPEG8% Control (no drug, light only)

70 60 50 40 30 20 10 FS (0 .3 FP )1 2% 0 J (0 .3 FP )1 8% 0 J (0 .3 )1 0 FS J (0 .3 FP )2 2% J (0 .3 FP )2 8% J (0 .3 FS )2 J (0 .0 FP 5) 2% 10 J (0 .0 FP 5) 8% 10 (0 J .0 5) 10 J Co nt ro l

0

Concentration of m-THPC in Foscan, FosPEG 2% and FosPEG 8% (mg/kg) Fig. 6. The percentage area of MC28 tumour necrosis induced after PDT with either 2 J or 10 J of light in female Hooded Lister rats at a drug-light interval (DLI) of 24 h, following intravenous injection of 0.3 and 0.05 mg kg−1 m-THPC in Foscan, FosPEG2%, and FosPEG 8% formulations. Negative control tumour tissue (received no drug or light). Data points show the mean ± SD, n = 4.

reached in a recent in vitro study using pegylated silica nanoparticles incorporating m-THPC [30]. All formulations were found to show peak m-THPC levels at the earliest time points (Fig. 2A–C) in the liver, spleen and lungs, which are all organs that constitute the RES. These organs are all involved in clearance pathways by acting as immunological filters of the blood [31]. The liver is known to contain ~ 20% of the total rat blood volume at any one time (~2.5 mL in 200 g rat) [32], which may explain these results. However, uptake of m-THPC in FosPEG 2% is almost half that of Foscan between 2 and 4 h in the liver and is completely cleared by 168 h (Fig. 2A), suggesting the PEG coating reduces protein binding by creating a more hydrophilic surface. A lower accumulation using FosPEG 8% is also observed in the liver compared to Foscan, but it is not significant. Interestingly, the biodistribution of m-THPC formulations in the lung tissue (Fig. 2C) correlates closely to the blood plasma pharmacokinetics (Fig. 1), as the uptake of m-THPC increases with liposomal pegylation. Following i.v. injection, the lungs are the first tissue to be perfused [27]. Combined with the high density of permeating blood vessels we speculate that circulating blood plasma may still have been present in the extracted tissue, resulting in a similar pattern of m-THPC accumulation. Complete clearance of Foscan and FosPEG 2% from the lungs by 168 h was observed; however a slower clearance rate was found for FosPEG8%. These slow distribution changes in the liver and lung could be the result of corresponding increased blood plasma levels of m-THPC in FosPEG 8%, as nanocarriers are redistributed to different tissues of the RES through macrophage uptake [33]. Quantitative Foscan studies with lung tissue have been investigated in vivo[27,34,35] and Fielding et al. [36] have demonstrated localisation of m-THPC in the macrophages of lung tissue at 72 h. To our knowledge, no studies have focused on the cumulative deposition of m-THPC-loaded pegylated liposomes in normal lung tissue, but findings such as these offer exciting potential in treating diseased or cancerous lung tissue with PDT. Other studies using comparable nanocarriers have indicated similar capabilities and applications [37,38]. The uptake for all m-THPC formulations is much lower in the kidneys in comparison to the major organs of the RES (Fig. 2D). Clearance using FosPEG 8% is in accordance with the pattern of excretion via the hepatobiliary pathway [9,39], as maximal renal filtration of 3–6 nm [16,40,41] prevents the passage of the pegylated liposomes (~120 nm). It is unclear however, as to why using Foscan and FosPEG 2% should display parallel pharmacokinetic profiles in the kidney, remaining at similar levels of accumulation over 168 h. The pharmacokinetic data for tumour tissue (Fig. 5A) display a peak in m-THPC accumulation with all formulations between 6 and 24 h following the administration of clinical doses of m-THPC (0.3 mg kg−1). This is in agreement with Foscan studies carried out by Jones et al. [26], which show an m-THPC peak around 18–24 h in LSBD1 tumours in the BDIX rat along with Cramers et al. in H-MESO1 tumour-bearing mice [42] after equal injected doses of m-THPC. More recently, Garrier et al. [25] found maximal Foscan uptake at 24 h in EMT6 tumour-bearing mouse models, comparable to Lassalle et al. [9] who observed similar concentrations at time periods between 6 and 15 h with Foslip in the same model and m-THPC dose (0.3 mg kg−1). However, in comparison to this study, approximately 90% less m-THPC accumulated in tumour tissue using Foslip than FosPEG (~0.1 μg g−1 vs. 1 μg g−1). Although it is not possible to directly compare results on account of differences in species and tumour models, it is apparent there are distinct differences in the tumour pharmacokinetics of m-THPC when delivered by liposomes with the addition of an inert PEG polymer coating. This has been well documented since the introduction of Stealth® liposomes [43–45]. The inherent accumulation of nanoparticles such as liposomes in tumour tissue can further be attributed to the enhanced permeability and retention (EPR) effect, since unlike normal tissue tumour vasculature is characterised by rapid angiogenesis [46], producing leaky blood vessels and a dysfunctional lymphatic drainage system [47]. This increase in tumour

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vessel permeability allows the nanocarriers to extravasate into the surrounding tumour tissue where they are retained due to poor clearance [11]. Combined with prolonged periods in blood circulation, pegylated liposomes are thought to accumulate passively in tumour tissue as a result of an increase in capillary fenestrae size (100–1200 nm) [48]. However, the dynamics are likely to be more complex since the mTHPC can also be slowly released from liposomes through binding to serum proteins [9,49,50]. In recent studies of silica nanoparticles incorporating non-covalently entrapped m-THPC, it was found that pegylation of the silica nanoparticle inhibited the release process due to steric hindrance at the surface, reducing interaction with serum proteins [30]. However, following cellular uptake, the liposomes are degraded to release the photosensitiser in its monomeric and photoactive form [9]. It is critical to examine the uptake and retention of m-THPC in skin tissue due to its well documented prolonged cutaneous photosensitivity post-PDT and subsequent complications with patient management. Although skin tissue was found to accumulate lower concentrations of m-THPC (≤ 0.1 μg g −1) compared to previous studies [9,26,51], surprisingly Foscan and FosPEG2% elicited similar kinetic profiles (Fig. 2F). However, the fast uptake (6 h) and increased accumulation of FosPEG 2% over Foscan in tumour tissue, which is illustrated in the skin to tumour tissue ratio (1.4 versus 6) (Fig. 5B), suggests this may be sufficient to reduce DLIs and drug doses without detriment to PDT efficacy and patient treatment. Elevated concentrations using FosPEG 8% were observed in the skin between 96 and 168 h, and for this reason these two time points were chosen to assess the effects of skin photosensitivity with sunlight (equivalent). Skin photosensitivity studies revealed extensive superficial damage to skin tissue with 0.3 mg kg–1 Foscan (i.v.) over the treatment area (5 mm2) after 30 min of light exposure at 100 mW cm–2 (sun light equivalent) with a DLI of 96 h. This is illustrated through histological examination by the thickening of the epidermis due to mass cell death at this site and an infiltration of inflammatory cells (indicated by arrows) (Fig. 4) indicative of severe erythema. The effects of light exposure (30 min) are slightly milder when using FosPEG 2% and 8% histologically, confirming damage to the epidermis at 96 h (Fig. 4). Visual assessment of skin effects at 96 h using FosPEG 8% (Fig. 3) correlates to chemical extraction results (Fig. 2F), whereby elevated accumulation of FosPEG 8% in skin tissue is reflected in the higher skin photodamage grading. However, the disparities between histological data are thought to be due to differences in the cellular localisation of m-THPC in skin tissue, which primarily determine the site and outcome of photoinduced damage [52]. Histologically there appears be less skin damage at a DLI of 168 h with all m-THPC formulations (Fig. 4). Foscan appears to elicit the most severe symptoms, in accordance with visual assessment (Fig. 3); however damage is more sporadic and occurs in isolated areas of the epidermis. No further effects are observed with FosPEG 2%. The skin histology for FosPEG 8% showed an infiltration of mast cells and neutrophils into the dermis (Fig. 4); however in one animal, widespread damage to the epidermis was detected (data not shown). This correlates to m-THPC accumulation for FosPEG 8% at 168 h obtained through chemical extraction but the irregularity probably accounts for the increase in photosensitive grading observed upon visual assessment. This is believed to be due to inter-animal variation, m-THPC localisation and small sampling size. All control skin samples including those that received no light (0 min exposure) and/or no m-THPC (data not shown) displayed no signs of cutaneous photosensitivity (Fig. 4). Recently, a reduction in skin photosensitivity has also been demonstrated using m-THPC encapsulation in micelles [53]. In PDT studies, tumour necrosis was measured as a percentage of whole tumour surface area from histological sections, as carried out previously [25]. At m-THPC doses of 0.3 mg kg−1 and exposure to 10 J of light, a high percentage of tumour necrosis was found for all formulations although no significant differences were evident, despite the greater mTHPC concentration in tumour tissue using the liposomal formulations

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(Fig. 5A). This may simply be due to the much larger light dose required to kill tissue at the tumour periphery, which may mask concentration differences, considering the complex dosimetry of interstitial light delivery [54,55]. A further factor is that the tumour periphery in sub-cutaneous models has been observed to be more resistant to PDT [55]. However, when either light energy or m-THPC doses are reduced to ≤20% of their original dose (2 J or 0.05 mg kg−1 m-THPC) a significant difference (p≤0.001) in percentage area tumour necrosis between Foscan versus FosPEG 2% and 8% is observed (Fig. 6) from ~30% necrosis to ≥40–50% necrosis. This implies not only that treatment times could be reduced, but also more importantly, that much lower doses of pegylated liposomal m-THPC may be administered in comparison to Foscan in order to elicit the same PDT effect witnessed at higher m-THPC doses. This pattern of PDT efficacy has also been observed in the treatment of arthritic joints [56] with pegylated liposomal m-THPC and in spontaneous squamous cell carcinomas in felines [57]. In our data set, PDT can be potentiated using the long circulating liposomes at 24 h, in accordance with both the tumour and plasma pharmacokinetics (Figs. 1, 5A and6). Although it is not feasible to discriminate clearly between damage to the vasculature and direct tumour cell damage mechanisms, it is likely that damage to the tumour vasculature, due to the increased levels of the FosPEG formulations in the circulation, makes an important contribution. We note also that histological examination showed signs of haemorrhagic damage. Tumour cell damage can then arise primarily through the accumulative effect of blood vessel collapse and hypoxia, with a lesser extent of direct tumour cell damage [42,58]. A similar mechanism was proposed by Lassalle et al. using a Foslip formulation. Although both FosPEG formulations are effective, at this stage it is unclear as to which FosPEG formulation is optimum for application in PDT. Moreover, the degree of surface pegylation is dependent upon establishing a balance between increasing the hydrophilicity of the liposome for macrophage evasion and maintaining the integrity of the lipid membrane to prevent destabilisation [14,16]. Further studies which focus on the cellular uptake and localisation of m-THPC would need to be performed. 5. Conclusion Our study has shown that m-THPC incorporation in pegylated liposomes can improve tumour selectivity in comparison to Foscan. An increase in blood plasma circulation, combined with the influence of the tumour microenvironment (EPR effect), are believed to contribute to enhanced tumour uptake of m-THPC in FosPEG liposomal formulations. Maximal tumour to skin ratios at ≤24 h with FosPEG 2% and 8% indicate that a shorter drug light interval could be adopted over current Foscan and treatment times could be reduced [9]. However, in a clinical setting, the main advantage of increased uptake and selectivity of m-THPC in pegylated liposomes into cancerous tissue is the reduction in damage to surrounding normal tissue, with less risk of scarring and functional damage. Enhancing the efficacy of PDT to tumours using lower administered drug doses could additionally cut therapeutic costs and simultaneously overcome adverse prolonged skin photosensitivity. Acknowledgements The research leading to these results has received funding from the European Community's Seventh Framework Programme (FP7/2007– 2013) under Grant Agreement No. 201031 (NANOPHOTO). This work was undertaken at UCLH/UCL which received a proportion of funding from the UK Department of Health's NIHR Biomedical Research Centres funding scheme. This work was also supported by the Experimental Cancer Medicine Centre, University College London. We should like to thank Biolitec (Jena, Germany) for supplying Foscan and pegylated liposomal m-THPC (FosPEG 2% and FosPEG 8%), Dr Sandy Mosse (NMLC, UCL) for technical assistance and expert pathologist Dr. Marco Novelli (Research Department of Pathology, UCL).

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