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Forest cover loss in the tropics is well known to cause warming at deforested sites, with maximum temperatures being particularly sensitive. Forest loss causes warming by altering local energy balance and surface roughness, local changes that can propagate across a wide range of spatial scales. Consequently, temperature increases result from not only changes in forest cover at a site, but also by the aggregate effects of non-local forest loss. We explored such non-local warming within Brazil’s Amazon and Cerrado biomes, the region with the world’s single largest amount of forest loss since 2000. Two datasets, one consisting of in-situ air temperature observations and a second, larger dataset consisting of ATs derived from remotely-sensed observations of land surface temperature, were used to quantify changes in maximum temperature due to forest cover loss at varying length-scales. We considered undisturbed forest locations (1 km ^2 in extent), and forest loss trends in annuli (‘halos’), located 1–2 km, 2–4 km, 4–10 km and 10–50 km from these undisturbed sites. Our research finds significant and substantial non-local warming, suggesting that historical estimates of warming due to forest cover loss under-estimate warming or mis-attribute warming to local change, where non-local changes also influence the pattern of temperature warming.

This letter uses satellite remote sensing to examine patterns of cropland expansion, cropland abandonment, and changing cropping frequency in Mato Grosso, Brazil from 2001 to 2011. During this period, Mato Grosso emerged as a globally important center of agricultural production. In 2001, 3.3 million hectares of mechanized agriculture were cultivated in Mato Grosso, of which 500 000 hectares had two commercial crops per growing season (double cropping). By 2011, Mato Grosso had 5.8 million hectares of mechanized agriculture, of which 2.9 million hectares were double cropped. We found these agricultural changes to be selective with respect to land attributes—significant differences ( p

Clinical proton beam quality assurance (QA) requires a simple and accurate method to measure the proton beam Bragg peak (BP) depth. Protoacoustics, the measurement of the pressure waves emitted by thermal expansion resulting from proton dose deposition, may be used to obtain the depth of the BP in a phantom by measuring the time-of-flight of the pressure wave. Rectangular and cylindrical phantoms of different materials (aluminum, lead, and polyethylene) were used for protoacoustic studies. Four different methods for analyzing the protoacoustic signals are compared. Data analysis shows that, for Methods 1 and 2, plastic phantoms have better accuracy than metallic ones because of the lower speed of sound. Method 3 does not require characterizing the speed of sound in the material, but it results in the largest error. Method 4 exhibits minimal error, less than 3 mm (with an uncertainty  ⩽1.5 mm) for all the materials and geometries. Psuedospectral wave-equation simulations (k-Wave MATLAB toolbox) are used to understand the origin of acoustic reflections within the phantom. The presented simulations and experiments show that protoacoustic measurements may provide a low cost and simple QA procedure for proton beam range verification as long as the proper phantoms and calculation methods are used.

Acoustic-based proton range verification (protoacoustics) is a potential in vivo technique for determining the Bragg peak position. Previous measurements and simulations have been restricted to homogeneous water tanks. Here, a CT-based simulation method is proposed and applied to a liver and prostate case to model the effects of tissue heterogeneity on the protoacoustic amplitude and time-of-flight range verification accuracy. For the liver case, posterior irradiation with a single proton pencil beam was simulated for detectors placed on the skin. In the prostate case, a transrectal probe measured the protoacoustic pressure generated by irradiation with five separate anterior proton beams. After calculating the proton beam dose deposition, each CT voxel's material properties were mapped based on Hounsfield Unit values, and thermoacoustically-generated acoustic wave propagation was simulated with the k-Wave MATLAB toolbox. By comparing the simulation results for the original liver CT to homogenized variants, the effects of heterogeneity were assessed. For the liver case, 1.4 cGy of dose at the Bragg peak generated 50 mPa of pressure (13 cm distal), a 2×  lower amplitude than simulated in a homogeneous water tank. Protoacoustic triangulation of the Bragg peak based on multiple detector measurements resulted in 0.4 mm accuracy for a δ-function proton pulse irradiation of the liver. For the prostate case, higher amplitudes are simulated (92-1004 mPa) for closer detectors (<8 cm). For four of the prostate beams, the protoacoustic range triangulation was accurate to  ⩽1.6 mm (δ-function proton pulse). Based on the results, application of protoacoustic range verification to heterogeneous tissue will result in decreased signal amplitudes relative to homogeneous water tank measurements, but accurate range verification is still expected to be possible.

A reply is provided to the points raised in the comment by Dr Sitek (2016 Phys. Med. Biol. 61 8941) on Polf et al (2015 Phys. Med. Biol. 60 7085).

The unique dose deposition of proton beams generates a distinctive thermoacoustic (protoacoustic) signal, which can be used to calculate the proton range. To identify the expected protoacoustic amplitude, frequency, and arrival time for different proton pulse characteristics encountered at hospital-based proton sources, the protoacoustic pressure emissions generated by 150 MeV, pencil-beam proton pulses were simulated in a homogeneous water medium. Proton pulses with Gaussian widths ranging up to 200 μs were considered. The protoacoustic amplitude, frequency, and time-of-flight (TOF) range accuracy were assessed. For TOF calculations, the acoustic pulse arrival time was determined based on multiple features of the wave. Based on the simulations, Gaussian proton pulses can be categorized as Dirac-delta-function-like (FWHM < 4 μs) and longer. For the δ-function-like irradiation, the protoacoustic spectrum peaks at 44.5 kHz and the systematic error in determining the Bragg peak range is <2.6 mm. For longer proton pulses, the spectrum shifts to lower frequencies, and the range calculation systematic error increases (⩽ 23 mm for FWHM of 56 μs). By mapping the protoacoustic peak arrival time to range with simulations, the residual error can be reduced. Using a proton pulse with FWHM = 2 μs results in a maximum signal-to-noise ratio per total dose. Simulations predict that a 300 nA, 150 MeV, FWHM = 4 μs Gaussian proton pulse (8.0 × 10(6) protons, 3.1 cGy dose at the Bragg peak) will generate a 146 mPa pressure wave at 5 cm beyond the Bragg peak. There is an angle dependent systematic error in the protoacoustic TOF range calculations. Placing detectors along the proton beam axis and beyond the Bragg peak minimizes this error. For clinical proton beams, protoacoustic detectors should be sensitive to <400 kHz (for -20 dB). Hospital-based synchrocyclotrons and cyclotrons are promising sources of proton pulses for generating clinically measurable protoacoustic emissions.

The purpose of this paper is to evaluate the ability of a prototype Compton camera (CC) to measure prompt gamma rays (PG) emitted during delivery of clinical proton pencil beams for prompt gamma imaging (PGI) as a means of providing in vivo verification of the delivered proton radiotherapy beams. A water phantom was irradiated with clinical 114 MeV and 150 MeV proton pencil beams. Up to 500 cGy of dose was delivered per irradiation using clinical beam currents. The prototype CC was placed 15 cm from the beam central axis and PGs from 0.2 MeV up to 6.5 MeV were measured during irradiation. From the measured data (2D) images of the PG emission were reconstructed. (1D) profiles were extracted from the PG images and compared to measured depth dose curves of the delivered proton pencil beams. The CC was able to measure PG emission during delivery of both 114 MeV and 150 MeV proton beams at clinical beam currents. 2D images of the PG emission were reconstructed for single 150 MeV proton pencil beams as well as for a 5   ×   5 cm mono-energetic layer of 114 MeV pencil beams. Shifts in the Bragg peak (BP) range were detectable on the 2D images. 1D profiles extracted from the PG images show that the distal falloff of the PG emission profile lined up well with the distal BP falloff. Shifts as small as 3 mm in the beam range could be detected from the 1D PG profiles with an accuracy of 1.5 mm or better. However, with the current CC prototype, a dose of 400 cGy was required to acquire adequate PG signal for 2D PG image reconstruction. It was possible to measure PG interactions with our prototype CC during delivery of proton pencil beams at clinical dose rates. Images of the PG emission could be reconstructed and shifts in the BP range were detectable. Therefore PGI with a CC for in vivo range verification during proton treatment delivery is feasible. However, improvements in the prototype CC detection efficiency and reconstruction algorithms are necessary to make it a clinically viable PGI system.

Recent studies have suggested that the characteristics of prompt gammas (PGs) emitted from excited nuclei during proton therapy are advantageous for determining beam range during treatment delivery. Since PGs are only emitted while the beam is on, the feasibility of using PGs for online treatment verification depends greatly on the design of highly efficient detectors. The purpose of this work is to characterize how PG detection changes as a function of distance from the patient as a means of guiding the design and usage of clinical PG imaging detectors. Using a Monte Carlo model (GEANT4.9.4) we studied the detection rate (PGs per incident proton) of a high purity germanium detector for both the total PG emission and the characteristic 6.13 MeV PG emission from (16)O emitted during proton irradiation. The PG detection rate was calculated as a function of distance from the isocenter of the proton treatment nozzle for: (1) a water phantom irradiated with a proton pencil beam and (2) a prostate patient irradiated with a scanning beam proton therapy treatment field (lateral field size: ∼6 cm × 6 cm, beam range: 23.5 cm). An analytical expression of the PG detection rate as a function of distance from isocenter, detector size, and proton beam energy was then developed. The detection rates were found to be 1.3 × 10(-6) for oxygen and 3.9 × 10(-4) for the total PG emission, respectively, with the detector placed 11 cm from isocenter for a 40 MeV pencil beam irradiating a water phantom. The total PG detection rate increased by ∼85 ± 3% for beam energies greater than 150 MeV. The detection rate was found to be approximately 2.1 × 10(-6) and 1.7 × 10(-3) for oxygen and total PG emission, respectively, during delivery of a single pencil beam during a scanning beam treatment for prostate cancer. The PG detection rate as a function of distance from isocenter during irradiation of a water phantom with a single proton pencil beam was described well by the model of a point source irradiating a cylindrical detector of a known diameter over the range of beam energies commonly used for proton therapy. For the patient studies, it was necessary to divide the point source equation by an exponential factor in order to correctly predict the falloff of the PG detection rate as a function of distance from isocenter.